Devices and Methods for Detection and Quantification of Immunological Proteins, Pathogenic and Microbial Agents and Cells

ABSTRACT

The present invention provides microfluidic pScreen™ devices for quantifying the concentration of DNA fragments in a liquid sample by using magnetic-responsive silica micro-beads and nonmagnetic-responsive silica micro-beads. The devices of the present invention allow for rapid, simple and inexpensive quantification of DNA fragment concentration in a sample. The devices do not require complex instrumentation and can be performed in less than three minutes. Moreover, they are compatible with complex samples including, without limitation, unpurified PCR amplification products, and thus can be expected to seamlessly integrate into various common molecular biology techniques and workflows.

CROSS REFERENCE TO RELATED APPLICATIONS

The present application is a continuation-in-part of U.S. patentapplication Ser. No. 15/666,918, filed Aug. 2, 2017, which is acontinuation of U.S. patent application Ser. No. 14/148,010, filed Jan.6, 2014, which is a continuation-in-part of U.S. patent application Ser.No. 13/862,899, filed Apr. 15, 2013, now U.S. Pat. No. 8,623,596, whichis a continuation application of U.S. patent application Ser. No.13/684,618, filed Nov. 26, 2012, now U.S. Pat. No. 8,445,192, issued May21, 2013, which is a continuation-in-part application of U.S. patentapplication Ser. No. 13/590,859, filed Aug. 21, 2012, which claimspriority to U.S. Provisional Patent Application No. 61/539,210, filedSep. 26, 2011, all of which are incorporated by reference herein intheir entirety.

FIELD OF THE INVENTION

The present invention relates to the field of immunoassay andmicrofluidic devices and, in particular, to a point-of-care diagnosticmethod and device for the detection and quantification ofmagnetic-responsive micro-beads conjugated with proteins, cells andmicrobial agents dispersed in a liquid sample.

BACKGROUND OF THE INVENTION

Current immunoassay technologies for the detection and quantification ofproteins rely on the specificity of the chemical interaction betweenantigens and antigen-specific antibodies. These tests may be classifiedinto two main groups: laboratory based-tests and point-of-care (POC)tests. Laboratory-based tests are sensitive and accurate, but require alaboratory setting and skilled technicians. POC tests are designed to beused in the field and require limited training, but they are far lesssensitive and accurate with most POC tests providing only binarypositive/negative or semi-quantitative results.

All current immunoassay technologies involve the formation of anantigen-antibody complex. The detection of the complex indicates thepresence of a targeted analyte in a sample. The antigen-antibody complexis detected by measuring the emission and/or reflection of light by thecomplex, when fluorescent-tagged antigen-specific antibodies areemployed, or in the case in which antibody-coated micro-beads are used,by measuring the emission and/or reflection of light, or the magneticmoment of the micro-beads forming the antigen-bead complex. In allcases, optical or magnetic detectors and electronic readers arerequired.

For example, the simplest, best known and widely used POC diagnosticassay is the lateral flow assay, also known as the immunochromatic test.In this test, the targeted analyte is bound to an analyte-specificantibody linked to latex or gold nanoparticles. The presence of theanalyte in the sample then is revealed by the formation of a visibleband, or line, which results from the agglutination or accumulation ofthe analyte-antibody-linked complex. The band typically is visiblemacroscopically to the naked eye. Devices to increase the assay'ssensitivity have been developed which can read color changes withmicroscopic sensitivity. Fluorescent or magnetic-labeled particles alsohave been used. In these cases, however, electronic readers to assesstest results are needed. Thus, although sensitivity of the assay mayincrease, the cost and complexity of the assay also increases.

In recent years, antibody-coated micro-beads have been increasingly usedfor the separation and detection of proteins. In the field ofimmunoassay diagnostics, the concentration of micro-beads is a proxy forthe concentration of targeted proteins in a sample. In theseapplications, it is necessary to identify the concentration ofmicro-beads in the sample solution. The micro-beads may be mademagnetically responsive by adding a magnetic core or layer to a polymerbead. The micro-beads then may be coated with a variety of molecules andproteins, referred to as ligands, which serve the purpose of binding thetargeted antigen via an antibody-antigen interaction. In addition,fluorescent dyes can be incorporated into the micro-beads making themoptically detectable. Recently, a diagnostic test for the proteintroponin using magnetic micro-beads has been proposed by Dittmer et al.(Philips Research Europe). In this assay, micro-beads coated withanti-troponin antibody are immobilized via antibody-troponin-antibodieson the surface of a micro-well with the aid of an applied magneticfield. The number of antibody-troponin-antibodies is measured byilluminating the bottom of the well and measuring the light reflected bythe immobilized micro-beads with an optical receiver. Methods for thedetection of E.coli also have been developed using immuno-magneticmicro-beads. In this case, the bacteria in the sample are measured bydetecting time-resolve fluorescence.

While micro-bead technology has matured in the last decade, thetechnology to quantify micro-bead concentration has lagged behind.Current methods include manual microscopes and automatic orsemi-automatic cell counters. Typically, micro-bead counting using amicroscope involves the manual, and often tedious, counting of beadsthrough a microscope objective. This method requires skilled techniciansin a laboratory setting, is time consuming and is subject to atechnician's interpretation. Cell counters require photo sensors todetect micro-beads automatically by measuring the light reflection of alaser beam hitting the micro-bead's surface. Cell counters, whileaccurate, are expensive and also require skilled technicians insophisticated laboratory settings. Lab-on-a-chip devices to detect andmeasure the concentration of protein-coated micro-bead concentrationalso have been developed. These devices, however, rely on traditionalapproaches, i.e., light reflection and detection using micro-scale lightand photo sensors and micro-scale magneto-resistance magnetometers.Thus, while greatly reducing the need for a laboratory setting andequipment, lab-on-a-chip devices still require electrical readers andtransducers. In addition, these devices typically include handset andconsumable components, resulting in increased manufacturing, calibrationand maintenance costs. Thus, these devices have limited applications inthe field of POC immunoassay diagnostics.

In addition to current immunoassay technologies which involve theformation of an antigen-antibody complex, quantification ofdeoxyribonucleic acid (DNA) is a fundamental step in many molecularbiology workflows. Whether quantifying isolated DNA fragments, amplifiedsegments from Polymerase Chain Reaction (PCR), or library and cloningcomponents, an accurate determination of the concentration of DNAfragments is important for experimental design and overall success.

Multiple methods to measure DNA concentration currently are available.These include optical methods, such as measuring the amount of UVradiation absorbed by the constituent nucleotides or the intensity offluorescence emission following addition of an intercalating dye, aswell as the comparison of an unknown sample against DNA standards ofknown concentration using gel electrophoresis or real-time PCR. Each ofthe methods have their advantages and disadvantages with regard tosensitivity, assay time, need of dedicated instrumentation/reagents, andsusceptibility to interference from contaminants.

Thus, there exists a need for a POC immunoassay device which has thesensitivity and specificity of laboratory-based immunoassay tests whilebeing simple to use and low cost, as well as for methods to detect andquantify magnetic-responsive micro-bead concentration in a samplespecimen. In addition, there exists a need for a POC device to measureDNA fragment concentration which has the sensitivity and specificity oflaboratory-based devices and methods for measuring DNA concentration.

SUMMARY OF THE INVENTION

The pScreen™ microfluidic immunoassay device, based on the inventionsdisclosed herein, fulfills all of the above-described needs in a singledevice. The detection and quantification of an unknown concentration ofanalyte in a liquid sample is obtained by exploiting the fluid-dynamicproperties of magnetic-responsive micro-beads in liquid solution ratherthan using optical effects or magnetic field sensing as in currenttechnologies. The unknown concentration of the target analyte is derivedby measuring the differential flow rate between the sample flow in twomicro-channels, one of which is under the influence of an appliedmagnetic field gradient. The present invention significantly reduces thecost and complexity of current laboratory-based immunoassay diagnostictests, and greatly increases one thousand-fold the sensitivity oflateral flow tests, while maintaining the specificity and accuracy oflaboratory-based tests, and the ability to detect targeted antigenconcentration over a predefined range.

In an embodiment of the present invention, there is provided a method ofdetecting and quantifying the concentration of magnetic-responsivemicro-beads in a fluid. The method comprises measuring flow rate ({tildeunder (O)}m) of a fluid in at least one test micro-channel (Cm) exposedto a magnetic field gradient with flow rate ({tilde under (O)}o) of thefluid in a calibration micro-channel (Co) not exposed to a magneticfield gradient, in which the micro-channels are kept at an equal andconstant pressure, calculating the ratio {tilde under (O)}m/{tilde under(O)}o, the difference {tilde under (O)}o−{tilde under (O)}m, and theratio ({tilde under (O)}o−{tilde under (O)}m)^(p)/({tilde under(O)}m)^(q), wherein p and q are derived through a calibration process,wherein the ratios {tilde under (O)}m/{tilde under (O)}o and ({tildeunder (O)}o−{tilde under (O)}m)^(p)/({tilde under (O)}m)^(q) are a proxyfor the number of magnetic-responsive micro-beads in the fluid, and thenquantifying the concentration of magnetic-responsive micro-beads in thefluid. The presence of magnetic-responsive micro-beads in the at leastone test micro-channel which is exposed to the magnetic field gradientcauses flocculation of the magnetic-responsive micro-beads in the fluidwhich reduces the flow rate of the fluid through the at least one testmicro-channel.

In another embodiment, there is provided a method for detecting andquantifying concentration of an analyte in a liquid sample. The methodcomprises adding a liquid sample to a liquid sample inlet of a reactionchamber. The reaction chamber has adsorbed on its surface a plurality ofimmobilized antigen-specific antibodies (Ab1) specific to an analyte.The surface of the reaction chamber also has a plurality ofmagnetic-responsive micro-beads desiccated thereon, in which each of theplurality of magnetic-responsive micro-beads is coated with anantigen-specific antibody (Ab2) specific to the analyte. The methodcomprises having the liquid sample incubate within the reaction chamber,which causes rehydration of the plurality of antibody-coatedmagnetic-responsive micro-beads as the liquid sample is added andagitated in the reaction chamber, which rehydration disperses theantibody-coated magnetic-responsive micro-beads in the liquid sample,binding the rehydrated antibody-coated magnetic-responsive micro-beadsas well as the antigen-specific antibodies immobilized on the surface ofthe reaction chamber to any analyte present in the liquid sample to formAb1-analyte-Ab2-coated magnetic-responsive micro-bead complexes on thesurface of the reaction chamber, having the liquid sample containing anyunbound antibody-coated magnetic-responsive micro-beads exit thereaction chamber through a chamber outlet and transfer through acontinuous fluid connection to a micro-channel splitter which bifurcatesto form a calibration micro-channel (Co) and at least one testmicro-channel (Cm). The at least one test micro-channel and thecalibration micro-channel are kept at an equal and constant pressure.The calibration micro-channel is in continuous fluid connection with agraduated column, and the at least one test micro-channel is incontinuous fluid connection with at least one graduated column. Each ofthe graduated columns has a graduated scale thereon. The methodcomprises measuring flow rate ({tilde under (O)}m) of the liquid samplein the at least one test micro-channel exposed to a magnetic fieldgradient with flow rate ({tilde under (O)}o) of the fluid in thecalibration micro-channel not exposed to a magnetic field gradient, inwhich the presence of any unbound antibody-coated magnetic-responsivemicro-beads in the at least one test micro-channel which is exposed tothe magnetic field gradient causes flocculation of the antibody-coatedmagnetic-responsive micro-beads in the liquid sample which reduces theflow rate of the liquid sample through the at least one testmicro-channel, calculating the ratio {tilde under (O)}m/{tilde under(O)}o, the difference {tilde under (O)}o−{tilde under (O)}m, and theratio ({tilde under (O)}o−{tilde under (O)}m)^(p)/({tilde under(O)}m)^(q), wherein p and q are derived through a calibration process,wherein the ratios {tilde under (O)}m/{tilde under (O)}o and ({tildeunder (O)}o−{tilde under (O)}m)^(p)/({tilde under (O)}m)^(q) are a proxyfor the number of magnetic-responsive micro-beads in the liquid sample,which is a proxy for the concentration of analyte in the liquid sample,and then quantifying the concentration of analyte in the liquid sample.

In another embodiment, there is provided a single use, portable,lab-on-card microfluidic pScreen™ magnetic-responsive micro-beadconcentration counter device for detecting and quantifying theconcentration of magnetic-responsive micro-beads in a liquid sample. Themicrofluidic device is comprised of a liquid sample inlet defined by anopening for accepting a liquid sample that contains a quantity ofmagnetic-responsive micro-beads. The liquid sample inlet is incontinuous fluid connection with a flow resistor, which is in continuousfluid connection with a micro-channel splitter which bifurcates to forma calibration micro-channel (Co) and at least one test micro-channel(Cm). The calibration micro-channel and the at least one testmicro-channel are kept at an equal and constant pressure. Thecalibration micro-channel is in continuous fluid connection with agraduated column, and the at least one test micro-channel is incontinuous fluid connection with at least one graduated column. The atleast one test micro-channel is exposed to a magnetic field gradient,which causes flocculation of the magnetic-responsive micro-beads in theat least one test micro-channel. The flocculation reduces the flow rate({tilde under (O)}m) of the liquid sample in the at least one testmicro-channel compared to the flow rate ({tilde under (O)}o) of theliquid sample in the calibration micro-channel. Each of the graduatedcolumns has a graduated scale thereon which provides a read-out of thetotal volume of the liquid sample collected in each of the graduatedcolumns, in which the total volume of the liquid sample collected in theat least one test micro-channel graduated column indicates theconcentration of magnetic-responsive micro-beads in the liquid sample.

In another embodiment of the invention, there is provided a single use,portable, lab-on-card microfluidic pScreen™ immunoassay device fordetecting and measuring an analyte in a liquid sample. The microfluidicpScreen™ immunoassay device comprises a liquid sample inlet defined byan opening for accepting the liquid sample. The liquid sample inlet isin continuous fluid connection with a flow resistor channel, which is incontinuous fluid connection with an assay inlet of a reaction chamber.The reaction chamber has adsorbed on its surface a plurality ofimmobilized antigen-specific antibodies (Ab1) specific to an analyte, aswell as having a plurality of magnetic-responsive micro-beads desiccatedthereon. Each of the plurality of magnetic-responsive micro-beads iscoated with an antigen-specific antibody (Ab2) specific to the analyte.Flow of the liquid sample through the reaction chamber rehydrates theplurality of antibody-coated magnetic-responsive micro-beads whichdisperses into the liquid sample. Any analyte present in the liquidsample binds to the dispersed antibody-coated magnetic-responsivemicro-beads as well as to the antigen-specific antibodies immobilized onthe surface of the reaction chamber to form Ab1-analyte-Ab2-coatedmagnetic-responsive micro-bead complexes. Any unbound antibody-coatedmagnetic-responsive micro-beads exit the reaction chamber through anassay outlet, which is in continuous fluid connection with amicro-channel splitter that bifurcates to form a calibrationmicro-channel (Co) and at least one test micro-channel (Cm), which arekept at an equal and constant pressure. The calibration micro-channel isin continuous fluid connection with a graduated column, and the at leastone test micro-channel in continuous fluid connection with at least onegraduated column. The at least one test micro-channel is exposed to amagnetic field gradient, which causes flocculation of themagnetic-responsive micro-beads in the at least one test micro-channel.The flocculation reduces the flow rate ({tilde under (O)}m) of theliquid sample in the at least one test micro-channel compared to theflow rate ({tilde under (O)}o) of the liquid sample in the calibrationmicro-channel. Each of the graduated columns has a graduated scalethereon which provides a read-out of the total sample volume collectedin each of the graduated columns, in which the total sample volumecollected in the at least one test micro-channel graduated columnindicates the concentration of analyte in the liquid sample.

In another embodiment, there is provided a method for detecting andquantifying concentration of an analyte in a liquid sample using asingle use, portable, lab-on-card microfluidic pScreen™ immunoassaydevice having a plurality of reaction chambers and a system of secondarymicro-channels. The method comprises adding a liquid sample to animmunoassay microfluidic device, the microfluidic device having a liquidsample inlet in continuous fluid connection with a liquid sample inletmanifold micro-channel that is in continuous fluid connection with aplurality of reaction chambers, each of the plurality of reactionchambers having adsorbed on its surface a plurality of immobilizedantigen-specific antibodies (Ab1) specific to an analyte each of theplurality of reaction chambers has adsorbed on its surface a pluralityof immobilized antigen-specific antibodies (Ab1) specific to an analyte,wherein the surface of each of the reaction chambers also has aplurality of magnetic-responsive micro-beads deposited thereon, each ofthe plurality of magnetic-responsive micro-beads coated with anantigen-specific antibody (Ab2) specific to the analyte; incubating theliquid sample within the plurality of reaction chambers, which causesrehydration of the plurality of antibody-coated magnetic-responsivemicro-beads, which rehydration disperses the antibody-coatedmagnetic-responsive micro-beads in the liquid sample; binding therehydrated antibody-coated magnetic-responsive micro-beads as well asthe antigen-specific antibodies immobilized on the surface of each ofthe reaction chambers to any analyte present in the liquid sample toform Ab1-analyte-Ab2-coated magnetic-responsive micro-bead complexes onthe surface of each of the reaction chambers; having the liquid samplecontaining any unbound antibody-coated magnetic-responsive micro-beadsexit the plurality of reaction chambers through outlet manifoldmicro-channels which is in continuous fluid connection with a connectormicro-channel, wherein one outlet manifold micro-channel also is incontinuous fluid connection with a delay micro-channel, the delaymicro-channel also in continuous fluid connection with the connectormicro-channel, wherein the connector micro-channel is in continuousfluid connection with a terminal flow splitter micro-channel thatbifurcates to form a calibration micro-channel (Co) and at least onetest micro-channel (Cm), wherein each of the outlet manifoldmicro-channels terminates in a passive valve, each of the passive valvessubstantially stopping, for a given period of time, the flow of theliquid sample from moving forward into the connector micro-channel;measuring the volume, Vm, of the liquid sample passing through the atleast one test micro-channel exposed to a magnetic field gradient, inwhich the presence of any unbound antibody-coated magnetic-responsivemicro-beads in the at least one test micro-channel which is exposed tothe magnetic field gradient causes flocculation of the antibody-coatedmagnetic-responsive micro-beads in the liquid sample which reduces theflow rate of the liquid sample through the at least one testmicro-channel, and the volume of the liquid sample passing through thecalibration micro-channel, Vo, not exposed to a magnetic field gradient;calculating the ratio Vm/Vo, the difference Vo−Vm, and the ratio(Vo−Vm)^(p)/(Vm)^(q), wherein p and q are derived through a calibrationprocess, wherein the ratios Vm/Vo and (Vo−Vm)^(p)/(Vm)^(q) are a proxyfor the number of magnetic-responsive micro-beads in the liquid sample,which is a proxy for the concentration of analyte in the liquid sample;and quantifying the concentration of analyte in the liquid sample.

In another embodiment of the invention, there is provided a single use,portable, lab-on-card microfluidic pScreen™ immunoassay device fordetecting and quantifying an analyte in a liquid sample having aplurality of reaction chambers and a system of secondary micro-channels.The microfluidic pScreen™ immunoassay device comprises a liquid sampleinlet, which is in continuous fluid connection with a liquid sampleinlet manifold micro-channel, which is in continuous fluid connectionwith a plurality of reaction chambers. Each of the reaction chambers hasan outlet manifold micro-channel where fluid exits the chambers. Each ofthe plurality of reaction chambers has adsorbed on its surface aplurality of immobilized antigen-specific antibodies (Ab1) specific toan analyte, and a plurality of magnetic-responsive micro-beads depositedthereon. Each of the plurality of magnetic-responsive micro-beads iscoated with an antigen-specific antibody (Ab2) specific to the analyte.When the liquid sample flows through the plurality of reaction chambers,the plurality of antibody-coated magnetic-responsive micro-beadsdisperses in the liquid sample and binds to any analyte present in theliquid sample. Any analyte present in the liquid sample also binds tothe antigen-specific antibodies immobilized on the surface of theplurality of reaction chambers to form Ab1 -analyte-Ab2-coatedmagnetic-responsive micro-bead complexes. Any unbound antibody-coatedmagnetic-responsive micro-beads exit each of the plurality of reactionchambers via the outlet manifold micro-channels. Each of the outletmanifold micro-channels terminates in a passive valve. The liquid sampleflow is substantially stopped for a period of time at each of thepassive valves, wherein the substantial stoppage of flow allows theanalyte to incubate with the Ab1 and Ab2 antibodies in the plurality ofreaction chambers. A primary flow splitter micro-channel is incontinuous fluid connection with one of the outlet manifoldmicro-channels, and the primary flow splitter micro-channel is incontinuous fluid connection with a delay micro-channel. The delaymicro-channel is in continuous fluid connection with a connectormicro-channel. The passive valves also are in continuous fluidconnection with the connector micro-channel. The connector micro-channelis in continuous fluid connection with a terminal flow splittermicro-channel which bifurcates to form a calibration micro-channel (Co)and at least one test micro-channel (Cm), which are kept at an equal andconstant pressure. The calibration micro-channel is in continuous fluidconnection with one or more calibration graduated columns, and the atleast one test micro-channel is in continuous fluid connection with oneor more test graduated columns. The test micro-channel is exposed to amagnetic field gradient, which causes flocculation of themagnetic-responsive micro-beads in the at least one test micro-channel.The flocculation reduces the flow rate ({tilde under (O)}m) of theliquid sample in the at least one test micro-channel compared to theflow rate ({tilde under (O)}o) of the liquid sample in the calibrationmicro-channel. The volume of liquid collected in the one or morecalibration graduated columns, Vo, and in the one or more test graduatedcolumns, Vm, are a proxy for the concentration of the analyte in theliquid sample.

Each of the reaction chambers, other than the reaction chamber that isin continuous fluid connection with the delay micro-channel, is incontinuous fluid connection with a secondary flow splittermicro-channel. Each secondary flow splitter micro-channel is incontinuous fluid connection with the outlet manifold micro-channel andwith an appendix micro-channel. Each of the appendix micro-channelsterminates in a vent port that is open to atmospheric pressure.

Each passive valve has three sharp edges and one continuous surfacecomprised of a sealing layer. The passive valves serve to substantiallystop the liquid sample flow for about 30 seconds to about 5 minutes dueto the three sharp edges which intersects with the connectormicro-channel. After this time, the passive valves burst sequentially sothat the liquid sample resumes flowing into the connector micro-channel.In contrast, the liquid sample flows freely through the delaymicro-channel and the appendix micro-channels via capillary action. Thisfree flow of liquid sample reduces the pressure gradient across thepassive valves enough so that the passive valves work to substantiallystop the flow of the liquid sample across the passive valves.

In accordance with the invention, the magnetic-responsive micro-beadsare deposited on the surface of each of the reaction chambers bydeposition of a micro-bead buffer solution containing the micro-beadsdispersed therein. The micro-bead buffer solution is comprised ofphosphate buffered saline. The micro-beads are deposited in nano-drops,in which each nano-drop has a volume of about 3 nl to 60 nl.

In accordance with the invention, mass density of the micro-bead buffersolution is increased to match mass the density of the micro-beads byadding additives to the micro-bead buffer solution. The additives mayinclude, without limitation, heavy water, glycerol, sucrose,polyethylene glycol, or a combination of two or more of the additives.

In accordance with the invention, all of the micro-channels of thedevice may be coated with a protein-free blocking solution which createsa hydrophilic film on the surface of the micro-channels, whichhydrophilic film decreases the liquid sample contact angle, increasesliquid sample flow rate, and decreases assay time.

In accordance with the invention, the liquid sample inlet can have aconical shape and a super-hydrophobic surface in order to create aconvex meniscus which creates pressure within the liquid sample that isgreater than the atmospheric pressure so that a pressure gradient iscreated which favors the flow of the liquid sample through the device.

In accordance with the invention, the micro-channel splitter of themicrofluidic device bifurcates to form one test micro-channel and onecalibration micro-channel. In another embodiment of the invention, themicro-channel splitter of the microfluidic device bifurcates to formthree test micro-channels and one calibration micro-channel, in whicheach of the three test micro-channels is in continuous fluid connectionwith one graduated column. In another embodiment of the invention, themicro-channel splitter of the microfluidic device bifurcates to formfour test micro-channels and one calibration micro-channel, in which thefour test micro-channels merge to be in continuous fluid connection withone graduated column.

Liquid samples that can be assayed in accordance with the embodiments ofthe invention include, without limitation, water, plasma, serum, buffersolution, urine, whole blood, blood analogs, and liquid solutions fromdilution of solid biological matter or other biological fluids.

Analytes that can be detected and quantified in accordance with theembodiments of the invention include, without limitation, proteins,protein fragments, antigens, antibodies, antibody fragments, peptides,RNA, RNA fragments, functionalized magnetic micro-beads specific toCD⁴⁺, CD⁸⁺ cells, malaria-infected red blood cells, cancer cells, cancerbiomarkers such as prostate specific antigen and other cancerbiomarkers, viruses, bacteria such as E. coli or other pathogenicagents.

The magnetic field gradient in accordance with the invention isgenerated from two magnets aligned lengthwise with the at least one testmicro-channel and along opposite poles to expose the at least one testmicro-channel to the magnetic field gradient. The at least one testmicro-channel is located between a gap formed between the opposite polesof the magnets. In another embodiment, the magnetic field gradient isgenerated by one magnet and a magnetic-responsive structure positionednear the at least one test micro-channel.

In accordance with the invention, the magnetic field generated can rangebetween about 0.05 Tesla (T) to about 0.5 T, and the magnetic fieldgradient that is generated can be about 10T/m or greater.

The total sample volume collected in the calibration micro-channelgraduated column serves as a control for parameters such as variation inviscosity between samples, level of hematocrit in blood samples,temperature and humidity fluctuations and sample volumes.

The present invention will be more fully understood from the followingdescription of the invention and by reference to the figures and claimsappended hereto.

In another embodiment, there is provided a microfluid device forquantifying the concentration of DNA fragments. The microfluid devicecomprises a sample inlet defined by an opening for accepting a liquidsample; a sealing layer; a reaction chamber; a micro-channel splitter; acalibration micro-channel; a test micro-channel; a control graduatedcolumn; a test graduated column; a venting hole located at the end ofeach of the graduated columns that is placed through the sealing layer;and a magnetic field gradient adjacent to the test micro-channel.

The liquid sample contains a mixture of magnetic-responsive silicamicro-beads, nonmagnetic-responsive silica micro-beads, and an aliquotof a DNA fragment sample, wherein the liquid sample is placed on theinlet of the microfluidic device.

The presence of DNA fragments triggers the formation of largemagnetic-responsive clusters, the magnetic-responsive clustersaggregating in proximity of the magnetic field gradient to produce alocalized restriction, the restriction retarding flow of the liquidsample in the test column.

The concentration of the nonmagnetic-responsive silica micro-beadsranges from between 10 μg/μl and 60 μg/μL in the liquid sample; and theconcentration of magnetic-responsive silica micro-beads ranges frombetween 1.0 μg/μl and 2.0 μg/μL in the liquid sample.

In an embodiment, the control graduated column and the test graduatedcolumn are connected to one another via a joining channel, and a ventinghole located at the end of the control column is placed through thesealing layer to stop the flow in the test channel when the liquidflowing in the control column fills to a desired volume.

In another embodiment, there is provided a microfluid passive valvelayout to stop the flow of liquid in a primary micro-channel when acontrol micro-channel fills to a desired volume. The microfluidicpassive valve layout comprises a primary micro-channel, a controlmicro-channel, a joining channel connecting the control and the primarymicro-channel, and a venting hole in the joining channel which is placedthrough the sealing layer, the venting hole placed such that the volumeof liquid filling the control micro-channel is the desired amount.

BRIEF DESCRIPTION OF THE DRAWINGS

A fuller understanding of the invention can be gained from the followingdescription when read in conjunction with the accompanying drawings inwhich:

FIG. 1 is a schematic illustration of the method for determining thenumber of magnetic-responsive micro-beads in a fluid, in which the ratiobetween the flow rate in the calibration micro-channel (Co) and the testmicro-channel (Cm) is measured, according to the embodiments of theinvention;

FIG. 2 is a schematic illustration of the microfluidic pScreen™magnetic-responsive micro-bead concentration counter device, having onetest micro-channel and one calibration micro-channel, according to theembodiments of the invention;

FIG. 3 is an artistic rendering of the microfluidic pScreen™ immunoassaydevice, according to the embodiments of the invention;

FIG. 4 is a schematic illustration of the microfluidic pScreen™magnetic-responsive micro-bead concentration counter device, havingthree test micro-channels and one calibration micro-channel, accordingto the embodiments of the invention;

FIG. 5 is a schematic illustration of the microfluidic pScreen™magnetic-responsive micro-bead concentration counter device, having fourtest micro-channels and one calibration micro-channel, according to theembodiments of the invention;

FIG. 6 is the schematic illustration of the method for determining theconcentration of analyte in a fluid, in which the sample analyte isbound to immobilized analyte-specific immobilized antibodies and toanalyte-specific coated magnetic-responsive micro-beads, the ratio{tilde under (O)}_(m)/{tilde under (O)}_(o), of the flow rate {tildeunder (O)}_(o) in the calibration micro-channel, Co, and the flow rate,{tilde under (O)}_(m), in the test micro-channel, Cm, is measured,according to the embodiments of the invention;

FIG. 7 is a schematic illustration of the pScreen™ immunoassay device,according to the embodiments of the invention;

FIGS. 8A, 8B and 8C are schematic illustrations of three different viewsof a reaction chamber of the pScreen™ immunoassay device, in which FIG.8A shows the secondary antibody (Ab2)-coated magnetic-responsivemicro-beads and primary antibody (Ab1)-capturing antibodies on thesurface of the reaction chamber; FIG. 8B shows theAb1-antigen-Ab2-magnetic-responsive micro-bead complexes immobilized onthe surface of the chamber; and FIG. 8C shows unbound, i.e., free,Ab2-magnetic-responsive micro-beads reaching the assay outlet of thereaction chamber, according to the embodiments of the invention;

FIG. 9 is a schematic illustration which shows the formation of theAb1-antigen-Ab2-magnetic-responsive micro-bead complexes as the samplewith the antigen flows through the reaction chamber and rehydrates themagnetic-responsive micro-beads, according to the embodiments of theinvention;

FIG. 10 is an artistic rendering of the pScreen™ immunoassay device,according to the embodiments of the invention;

FIG. 11 is a graph showing reduction in fluid flow rate , i.e., theratio between flow rate in the test (with magnetic field) andcalibration (without magnetic field) micro-channels versus the number ofmagnetic-responsive micro-beads in the flocculation region;

FIG. 12 is a photomicrograph showing magnetically-induced flocculationof magnetic-responsive micro-beads at a concentration of about 2,000micro-beads/μl in a test micro-channel, according to the embodiments ofthe invention; the inset shows the variation in flocculation at 300seconds, 600 seconds and 800 seconds;

FIG. 13 is a graph showing the concentration of magnetic-responsivemicro-beads in a solution obtained using the pScreen™ device versus thenominal concentration as tested by standard hemocytometry;

FIG. 14 is a graph showing the concentration of immunoglobulin (IgG) ina solution obtained using the pScreen™ immunoassay;

FIG. 15 is a graph showing a calibration curve describing therelationship between (Vo−Vm)/(Vm)² and the concentration ofmagnetic-responsive micro-beads, according to the embodiments of theinvention;

FIG. 16A and 16B are schematic illustrations of the pScreen™ immunoassaydevice, FIG. 16A showing a top view, and FIG. 16B showing a side view,in accordance with the invention;

FIG. 17A and 17B and 17C are schematic illustrations of the system ofpassive valves, with FIG. 17A and 17B showing a top view, and FIG. 17Cshowing a perspective view of the intersection of one outlet manifoldmicro-channel with the connector micro-channel, in accordance with theinvention;

FIG. 18 is a schematic illustration of the pScreen™ immunoassay device,in accordance with the invention;

FIG. 19A is a schematic illustration of the pScreen™ immunoassay deviceshowing the sample inlet and the device's longitudinal axis. FIG. 19B isa schematic illustration of a cross-section of the pScreen™ immunoassaydevice along the longitudinal axis at B-B, showing the two levels of thedevice, in accordance with the invention;

FIG. 20A and 20B are schematic illustrations of the pScreen™ immunoassaydevice inlet, with FIG. 20A showing a cross-section of the inlet, andFIG. 20B showing an expanded view of the inlet contact angle and inletcone angle, in accordance with the invention;

FIG. 21 is a schematic illustration of one reaction chamber, one outletmicro-channel manifold, and a delay micro-channel. Ro represents thefluid resistance of the sum of the reaction chamber and outlet manifoldmicro-channel, and Rl represents the fluid resistance in the delaymicro-channel, in accordance with the invention;

FIG. 22 is a schematic illustration of one reaction chamber, one outletmicro-channel manifold, and one appendix micro-channel. Ro representsthe fluid resistance of the sum of the reaction chamber and outletmanifold micro-channel, Rl represents the fluid resistance in theappendix micro-channel, Po represents the pressure at the flow splitter,Pv represents the pressure in the fluid at the fluid meniscus, Parepresents atmospheric pressure at the fluid meniscus, and Pl representsthe pressure inside the fluid meniscus at the appendix micro-channel, inaccordance with the invention;

FIG. 23 is a schematic top view illustration of one reaction chambershowing an array of micro-bead drops, in accordance with the invention;

FIG. 24 is a plot of the number of micro-beads per drop deposited on areaction chamber versus dispensing time. The vertical axis is theaverage number of micro-beads per drop as a percentage of the number ofmicro-beads per drop in the first drop deposited at time t=0, and thehorizontal axis is the dispensing time;

FIG. 25A is a top view schematic illustration of the pScreen™microfluidic device for measuring DNA concentration; and FIG. 25B is aside view schematic illustration of the pScreen™ microfluidic device formeasuring DNA concentration, according to the embodiments of theinvention;

FIG. 26A is a top view schematic illustration of the pScreen™microfluidic device for measuring DNA concentration which includes ajoining channel; and FIG. 26B is a side view schematic illustration ofthe pScreen™ microfluidic device for measuring DNA concentration whichincludes a joining channel, according to the embodiments of theinvention;

FIG. 27 is a plot of Delta-L (mm) vs DNA Fragment Scaled Concentrationobtained for various DNA Fragments lengths from 150 Base Pairs (BPs) toabove 10,000 and including genomic DNA; and

FIG. 28 is a plot of the Concentration of the DNA Fragments (ng/μl)obtained with the p-Screen Device vs. the True Concentration of DNAFragments.

DETAILED DESCRIPTION OF THE INVENTION

As used herein, the terms “analyte” and “antigen” are meant to beinterchangeable.

As used herein, the terms “calibration micro-channel(s)” and “controlmicro-channel(s)” are meant to be interchangeable.

As used herein, the terms “magnetic-responsive micro-beads,” and“magnetic micro-beads” are meant to be interchangeable.

As used herein, the terms “microfluidic device” and “microfluidic card”are meant to be interchangeable.

The present invention provides a flow rate-based method for detectingand quantifying the concentration, i.e., number, of magnetic-responsivemicro-beads in a fluid. The ratio, {tilde under (O)}m/{tilde under(O)}o, between the flow rate ({tilde under (O)}m) in a testmicro-channel (Cm) exposed to a localized high-gradient magnetic field,and the unperturbed flow rate ({tilde under (O)}o) in a calibration, orcontrol, micro-channel (Co) not exposed to the localized high-gradientmagnetic field, is a monotonic function of the number ofmagnetic-responsive micro-beads flowing through the test micro-channel.That is:

{tilde under (O)}m(N _(m))/{tilde under (O)}o=f(N _(m))   Equation (1)

where Nm is the total number of magnetic-responsive micro-beadstransported by the fluid into the localized high-magnetic field region.Both micro-channels are held at an equal and constant pressure. As shownin FIG. 1, a fluid 12 seeded with magnetic-responsive micro-beads 15flows into two micro-channel inlets 14, 14′ and out of two micro-channeloutlets 16, 16′. The test micro-channel 22 on the right is exposed to ahigh-gradient magnetic field generated by two magnets 24, 24′. Themagnets are positioned as shown in FIG. 1 and in the inset. In anembodiment, the magnetic field gradient is generated by one magnet and amagnetic-responsive structure (not shown) positioned near the testmicro-channel 22. A magnetic-responsive structure may be made of ametallic material with ferromagnetic, super-paramagnetic or paramagneticproperties, such that upon application of an external magnetic field themagnetic-responsive structure generates an induced magnetic field. Thestructure is geometrically shaped, e.g., cylindrically-shaped, in orderto generate a magnetic field gradient in the region occupied by the testmicro-channel. Equation 1 applies to a wide range of magnetic-responsivemicro-bead concentrations, ranging from about 50 micro-beads/μl to about2×10⁶ micro-beads/μl. The upper and lower limits, however, are afunction of the micro-channels' size and magnetic field topology. Hence,both upper and lower limits may vary based on these parameters.

Because f(Nm) is a monotonic function of Nm, it also holds that:

N _(m) =ƒ ⁻¹({tilde under (O)}m(N _(m))/{tilde under (O)}o)   Equation(2)

Thus, according to Equation 2, the number of magnetic-responsivemicro-beads in a given fluid is a monotonic function of the ratio {tildeunder (O)}m/{tilde under (O)}o. Thus, the number of magnetic-responsivemicro-beads can be determined by measuring the ratio {tilde under(O)}m/{tilde under (O)}o in the two micro-channels, configured as shownin FIG. 1. In other words, the ratio {tilde under (O)}m/{tilde under(O)}o is a specific proxy for the number of magnetic-responsivemicro-beads in a given fluid.

The analytical form of the function depends on the geometry, i.e.,length and inner diameter of the two micro-channels, magnetic fieldtopology, and the size of the magnetic-responsive micro-beads. Inaddition, the difference {tilde under (O)}_(o)−{tilde under (O)}_(m),and the ratio ({tilde under (O)}_(o)−{tilde under (O)}_(m))^(p)/({tildeunder (O)}_(m))^(q), where p and q are derived through a calibrationprocess, are a proxy for the number of magnetic-responsive micro-beadsin the fluid. The parameters p and q are obtained as followed. Asolution containing a known concentration of micro-beads and of knownvolume is passed through the micro-channels and the flow rate {tildeunder (O)}m and {tilde under (O)}o are measured. Then, a solutioncontaining the same concentration of magnetic-responsive micro-beads butof larger volume similarly is passed through the micro-channels. Thisprocess is repeated several times. Then, the ratio ({tilde under(O)}o−{tilde under (O)}m)^(p)/({tilde under (O)}m)^(q), with p and q setequal to 1, are plotted versus the volume of each sample. Using thewell-known least square regression method, p and q are determined byenforcing the condition that the ratios ({tilde under (O)}o−{tilde under(O)}m)^(p)/({tilde under (O)}m)^(q) versus sample volume form ahorizontal straight line with slope equal to zero.

The present invention further provides a microfluidic pScreen™magnetic-responsive micro-bead concentration counter device fordetecting and quantifying magnetic-responsive micro-bead concentrationin a liquid sample. This device leverages the previously described flowrate-based detection and quantification method.

FIGS. 2 and 3 show the pScreen™ microfluidic device 5 for the detectionand quantification of magnetic-responsive micro-beads in a liquid sampleof the present invention. The microfluidic device 5 comprises a liquidsample inlet 8 in which a liquid sample, or specimen, which contains anunknown amount of magnetic-responsive micro-beads, is applied. From theliquid sample inlet 8, the liquid sample, self-propelled by capillaryaction, flows through a flow resistor 32 (shown in FIG. 3) and enters amicro-channel splitter 18. The micro-channel splitter 18 bifurcates intotwo smaller micro-channels: a calibration micro-channel 20 and a testmicro-channel 22. The two micro-channels 20, 22 are identical in lengthand inner diameter (best shown in FIG. 2).

The concentration of magnetic-responsive micro-beads that can bedetected and quantified using the methods and devices of the inventionis about 50 micro-beads/μl to about 2×10⁶ micro-beads/μl; and thediameter of the magnetic-responsive micro-beads is about 0.2 μm to about20 μm. In an embodiment, the diameter of the magnetic micro-beads isabout 4.0 μm.

In accordance with the invention, the test micro-channel and thecalibration micro-channel are made of a capillary tube, in which thelength of the capillary tube is about 0.2 cm to about 20 cm. In anembodiment, the length of the capillary tube is about 3.0 cm to about7.5 cm. In another embodiment, the length of the capillary tube is about1.5 cm.

In an embodiment, the length of the calibration micro-channel 20 and thetest micro-channel 22 is about 0.2 cm to about 20 cm. In anotherembodiment, the length of the two micro-channels 20, 22 is about 3.0 cmto about 7.5 cm. In still another embodiment, the length of the twomicro-channels 20, 22 is about 1.5 cm.

In an embodiment, the inner diameter of the calibration micro-channel 20and the test micro-channel 22 is about 50 μm to about 500 μm. In anotherembodiment, the inner diameter of the two micro-channels 20, 22 is about50 μm.

A magnetic field gradient is applied only to the test micro-channel 22.The magnetic field gradient is generated by small rare-earth (e.g.,neodymium) permanent magnet and ferromagnetic (e.g., nickel, iron) polestructures (not shown) which serve as a magnetic concentrator 54 (shownin FIG. 3) specifically designed to concentrate the magnetic field,hence creating a high magnetic field gradient (of about 100 T/m). In thecalibration channel 20, the liquid sample flows freely at a very lowvelocity (Reynolds number around 1) and shear rate range (1 to 400 s⁻¹).In the test micro-channel 22, the magnetic field gradient inducesmicro-bead flocculation if magnetic-responsive micro-beads are presentin the sample. Even in very small concentrations, as low as <50micro-beads/μ1, the flow rate through the test micro-channel 22 will bereduced due to the formation of the magnetically-induced micro-beadflocculation. After flowing through the calibration and testmicro-channels 20, 22, a volume of liquid sample is collected in twograduated columns 26, 26′, both of equal size and volume.

Each graduated column 26, 26′ has a graduated scale thereon 30 whichprovides an easy to interpret read-out system of the total sample volumecollected in each graduated column 26, 26′. The graduated columns' 26,26′ length and cross section, as well as the respective scales 30thereon, are designed to be visible to the naked eye. Unlike current POCread-out devices, the read-out system of the microfluidic device of thepresent invention does not require electrical transducers and/orsensors. As shown in FIGS. 2 and 3, both graduated columns 26, 26′ areclearly visible. As shown in FIG. 3, the microfluidic device 5 may beconfigured in a cartridge 58 which fits into a holder 56. The read-outobtained by the microfluidic device 5 can be determined at any time bydirect comparison of the fluid in the two graduated columns 26, 26′.

Referring now to FIG. 2, the micro-channel configuration provides forthe concentration of micro-beads to be a monotonic function only of thevolumes Vo and Vm (Vo and Vm are the volumes collected at themicro-channel outlets 16, 16′ of the calibration micro-channel 20 (Co)and test micro-channel 22 (Cm), respectively; shown in FIG. 1). Thisapproach allows the user to read out the result provided by the pScreen™microfluidic device of the present invention at any time while the assayis running or at any time after the assay has been completed.

Given the relationship in Equation (1), and because the flow rate in thecalibration micro-channel 20 is constant and the magnetic-responsivemicro-beads are uniformly distributed in the sample fluid, and bydefinition ρ=dN/dV and dV={tilde under (O)}dt, where N is the number ofmicro-beads, {tilde under (O)} the flow rate, and V the fluid volume,the below equations are satisfied at any time instances:

$\begin{matrix}{{t = {{\int\limits_{0}^{No}\frac{{dN}^{\prime}}{\rho \; Q_{0}}} = \frac{N_{0}}{\rho \; Q_{0}}}},} & {{Equation}\mspace{11mu} (3)} \\{t = {\int\limits_{0}^{Nm}{\frac{{dN}^{\prime}}{\rho \; Q_{m}}.}}} & {{Equation}\mspace{14mu} (4)}\end{matrix}$

where, t is the time, ρ is the magnetic-responsive micro-beadconcentration in the sample specimen 12, {tilde under (O)}o and {tildeunder (O)}m are the flow rates in the calibration and testmicro-channels 20, 22, respectively, No and Nm is the number ofmagnetic-responsive micro-beads passing through the calibration and testmicro-channels 20, 22, respectively, and the prime symbol inside theintegral, dN′, indicates, according to standard convention, that theintegral operation is computed on N variable

It thus follows that:

$\begin{matrix}{{{N_{0} = {g( N_{m} )}},{{with}\text{:}}}{{{g(N)} \equiv}\;,{{{and}\text{:}\mspace{11mu} } = {{\frac{Q_{m}}{Q_{0}}.{Thus}}\text{:}}}}} & {{Equation}\mspace{14mu} (5)} \\{{N_{m} = {{g^{- 1}( N_{0} )}.{Since}}},} & {{Equation}\mspace{14mu} (6)} \\{{{N_{0} = {\rho \; V_{0}}},{{{and}\mspace{14mu} N_{m}} = {\rho \; V_{m}}},{{we}\mspace{14mu} {have}\mspace{14mu} {that}\text{:}}}{{\rho \; V_{m}} = {g^{- 1}( {\rho \; V_{0}} )}}} & {{Equation}\mspace{14mu} (7)} \\{{{Hence}\text{:}\mspace{14mu} \rho} = {{F( {V_{0},V_{m}} )}.}} & {{Equation}\mspace{14mu} (8)}\end{matrix}$

Thus, the pScreen™ microfluidic device of the present invention providesa comparative read-out system in which the magnetic-responsivemicro-bead concentration, ρ, is a monotonic function of only Vm, thevolume flowing through the test micro-channel 22 where themagnetic-induced flocculation forms and Vo, the volume flowing throughthe calibration micro-channel 20 without the magnetic-inducedflocculation.

The comparative read-out system of the pScreen™ microfluidic device ofthe present invention greatly simplifies the detection andquantification of magnetic-responsive micro-bead concentration in aliquid sample. In addition, this comparative read-out system has thesignificant advantage of virtually eliminating common-mode error (withthe calibration graduated column 26 acting as a control), such asvariation in viscosity between samples, level of hematocrit in bloodsamples, temperature and humidity fluctuation of the test environment,and sample volume. The pScreen™ microfluidic device of the presentinvention thus provides a stand-alone device for the detection andquantification of magnetic-responsive micro-bead concentration in liquidsamples over a wide range of concentrations and micro-bead sizes.

FIG. 4 shows an alternate embodiment of the microfluidic device 5 of theinvention. In this embodiment, the test micro-channel (Cm) is split intothree test micro-channels 22, 22′, 22″ which run parallel to one other.The three test micro-channels 22, 22′ and 22″ are of the same length buthave a different inner diameter from one another. Each testmicro-channel 22, 22′, 22″ is connected to a separate graduated column26. In an embodiment, the first test micro-channel 22 has an innerdiameter of about 50 μm to about 500 μm, the second test micro-channel22′ has an inner diameter of about 100 μm to about 250 μm, and the thirdtest micro-channel 22″ has an inner diameter of about 250 μm to about 5mm. In another embodiment, the first test micro-channel 22 has an innerdiameter of about 50 μm, the second test micro-channel 22′ has an innerdiameter of about 100 μm, and the third test micro-channel 22″ has aninner diameter of about 250 μm. The inner diameter of the calibrationmicro-channel 20 is such that the area of the cross-section of thecalibration micro-channel 20 is identical to the sum of the areas of thecross-sections of the three test micro-channels 22, 22′, 22″.

For a given amount of magnetic-responsive micro-beads entering each ofthe three test micro-channels 22, 22′ 22″, the third, largest testmicro-channel 22″ experiences the lowest reduction in flow rate, thesecond, middle-sized test micro-channel 22′ experiences a reduction inflow rate greater than in the third, largest test micro-channel 22″, andthe first, smallest test micro-channel 22 experiences the greatestreduction in flow rate. In addition, the first, smallest testmicro-channel 22 will tend to clog before the second, middle-sized testmicro-channel 22′ and the third, largest test micro-channel 22″, and themiddle-sized test micro-channel 22′ will tend to clog before the largesttest micro-channel 22″. Hence, the device in accordance with thisembodiment allows measurement of a wide range of concentrations ofmagnetic-responsive micro-beads, in which the first, smallest testmicro-channel 22 allows for finely-tuned measurements ofmagnetic-responsive micro-beads at low concentrations and the third,largest test micro-channel 22″ allows for gross measurements ofmagnetic-responsive micro-beads at high concentrations.

FIG. 5 shows an additional alternate embodiment of the invention. Inthis embodiment, the test micro-channel 22 (Cm) is split into fourmicro-channels which run parallel to each other. The four testmicro-channels 22 have the same length and inner diameter. In anembodiment, each of the four test micro-channels has an inner diameterof about 12.5 μm to about 125 μm. In another embodiment, each of thefour test micro-channels has an inner diameter of about 12.5 μm. Theinner diameter of the calibration micro-channel 20 is such that the areaof the cross-section of the calibration micro-channel 20 is identical tothe sum of the areas of the cross-sections of the four testmicro-channels 22, 22′, 22″.

The four test micro-channels 22 rejoin to connect to one graduatedcolumn 26. If no magnetic-responsive micro-beads flow into the fourtest-micro-channels 22 and the one calibration micro-channel 20, thenthe flow rate of the fluid through the calibration micro-channel 20 isthe sum of the flow rates in each of the test micro-channels. Equations(1) through (8) also apply in this embodiment, however, because thereare four parallel test micro-channels compared to one testmicro-channel, a greater volume of fluid can flow through the device ina shorter amount of time, thus allowing a user to obtain a read out ofresults of the pScreen™ microfluidic device in a shorter period of time.

The present invention also provides a flow rate-based method fordetecting and quantifying concentration of an analyte in a liquidsample. The analyte can include, without limitation, proteins, proteinfragments, antigens, antibodies, antibody fragments, peptides, RNA, RNAfragments, cells, cancer cells, viruses, and other pathogenic agents.

The method according to this embodiment comprises adding a liquid sampleto a liquid sample inlet of a reaction chamber. The reaction chamber hasadsorbed on its surface a plurality of immobilized antigen-specificantibodies (Ab1) specific to an analyte. The surface of the reactionchamber also has a plurality of magnetic-responsive micro-beadsdesiccated thereon, in which each of the plurality ofmagnetic-responsive micro-beads is coated with an antigen-specificantibody (Ab2) specific to the analyte. The method comprises having theliquid sample incubate inside the reaction chamber, which causesrehydration of the plurality of antibody-coated magnetic-responsivemicro-beads as the liquid sample is added and agitated in the reactionchamber, which rehydration disperses the antibody-coatedmagnetic-responsive micro-beads in the liquid sample, binding therehydrated antibody-coated magnetic-responsive micro-beads as well asthe antigen-specific antibodies immobilized on the surface of thereaction chamber to any analyte present in the liquid sample to formAb1-analyte-Ab2-coated magnetic micro-bead complexes on the surface ofthe reaction chamber, having the liquid sample containing any unboundantibody-coated magnetic-responsive micro-beads exit the reactionchamber through a chamber outlet and transfer through a continuous fluidconnection to a micro-channel splitter which bifurcates to form acalibration micro-channel (Co) and at least one test micro-channel (Cm).The at least one test micro-channel and the calibration micro-channelare kept at an equal and constant pressure. The calibrationmicro-channel is in continuous fluid connection with a graduated column,and the at least one test micro-channel is in continuous fluidconnection with at least one graduated column. Each of the graduatedcolumns has a graduated scale thereon. The method comprises measuringflow rate ({tilde under (O)}m) of the liquid sample in the at least onetest micro-channel exposed to a magnetic field gradient with flow rate({tilde under (O)}o) of the fluid in the calibration micro-channel notexposed to a magnetic field gradient, in which the presence of anyunbound antibody-coated magnetic-responsive micro-beads in the at leastone test micro-channel which is exposed to the magnetic field gradientcauses flocculation of the antibody-coated magnetic-responsivemicro-beads in the liquid sample which reduces the flow rate of theliquid sample through the at least one test micro-channel, andcalculating the ratio {tilde under (O)}m/{tilde under (O)}o, thedifference {tilde under (O)}o−{tilde under (O)}m, and the ratio ({tildeunder (O)}o−{tilde under (O)}m)^(p)/({tilde under (O)}m)^(q), wherein pand q are derived through a calibration process, and wherein the ratios{tilde under (O)}m/{tilde under (O)}o and ({tilde under (O)}o−{tildeunder (O)}m)^(p)/({tilde under (O)}m)^(q) are a proxy for the number ofmagnetic-responsive micro-beads in the liquid sample, which is a proxyfor the concentration of analyte in the liquid sample.

As shown in FIG. 6, a liquid sample 12 is added, Step 1, to a reactionchamber 34, which has adsorbed on its surface a plurality of immobilizedantigen-specific antibodies (Ab1) (not shown) and contains a pluralityof magnetic-responsive micro-beads coated with antigen-specificantibodies (Ab2) 50 desiccated on the surface of the reaction chamber34. In Step 1, by adding the liquid sample to the reaction chamber 34,the antibody-coated magnetic-responsive micro-beads 50 are rehydratedand they, as well as the antigen-specific antibodies immobilized on thesurface of the reaction chamber 34, bind to any analyte present in theliquid sample 12 to form Ab1-analyte-Ab2-coated magnetic-responsivemicro-bead complexes (not shown) on the surface of the reaction chamber34, with any unbound magnetic-responsive micro-beads 50 free to flow(Step 2) into two micro-channel inlets 14, 14′ and out of twomicro-channel outlets 16, 16′. The test micro-channel 22 on the right isexposed to a high-gradient magnetic field generated by two magnets 24.As described in the previous paragraph, the number ofmagnetic-responsive micro-beads in a liquid sample passing through twomicro-channels, a test micro-channel (Co) and a calibration, or controlmicro-channel (Cm), is proportional to the ratio, {tilde under(O)}_(m)/{tilde under (O)}o, between the flow rate ({tilde under (O)}m)in a micro-channel (Cm) exposed to a localized high-gradient magneticfield, and the unperturbed flow rate ({tilde under (O)}o) in amicro-channel (Co) not exposed to the localized high-gradient magneticfield. Therefore, by measuring the flow rates {tilde under (O)}m and{tilde under (O)}o, the concentration of analyte in the liquid samplecan be determined. The method applies to a wide range of antigenconcentration, from about 0.01 ng/ml to about 1.0 μg/ml.

The present invention further provides a pScreen™ microfluidicimmunoassay device for the detection and quantification of proteins,protein fragments, antigens, antibodies, antibody fragments, RNA, RNAfragments, cells, cancer cells, viruses, and other pathogenic agents.This device leverages the previously described method for detecting andquantifying concentration of an analyte in a liquid sample.

Principle of Operation

In one embodiment of the invention, as shown in FIG. 7, the pScreen™microfluidic immunoassay device 10 has a liquid sample inlet 8, a flowresistor channel 32, a reaction chamber 34, a micro-channel splitter 18,a test micro-channel 22, a calibration micro-channel 20 and graduatedreadout columns 26, 26′. In use, a liquid sample, or specimen, which maycontain an unknown amount of a target analyte, is applied into theliquid sample inlet 8. By capillary action, the sample is self-propelledand transferred from the liquid sample inlet 8 into the reaction chamber34 via the flow resistor channel 32. The flow rate of the liquid sampleis determined by the cross-section and length of the flow resistorchannel 32 and the surface tension of the device material and sampleliquid, as well as by the binding kinetic reaction between the analyte,i.e., antigen, and antibody in the reaction chamber 34.

As shown in FIG. 8A, the reaction chamber 34 is coated withantigen-specific antibodies (Ab1) 46 immobilized onto the surface of thereaction chamber 34. The antigen-specific antibodies (Ab1) 46, referredto as capturing antibodies, may be primary or secondary antibodies. Theantigen-specific antibodies (Ab1) 46 are bound to the surface of thereaction chamber 34 via adsorption. The surface of the reaction chamber34 also is coated with antibody-coated magnetic-responsive micro-beads50 by desiccation. The magnetic-responsive micro-beads may be desiccatedon the same region of the device where Ab1 antibodies 46 are bound, orin a region preceding where the Ab1 antibodies 46 are bound. Theantibody-coated magnetic-responsive micro-beads 50 are coated withantigen-specific antibodies (Ab2). These antigen-specific antibodies(Ab2) may be primary or secondary antibodies. As the liquid sample flowsinto the reaction chamber 34 via a chamber inlet 36, theantigen-specific antibody-coated magnetic-responsive micro-beads 50 arerehydrated and dispersed in the liquid, and any antigen molecules 48contained in the sample bind to the antigen-specific antibodies (Ab1) 46immobilized on the surface of the reaction chamber, and to theantigen-specific antibodies (Ab2) coating the magnetic micro-beads,forming Ab1-antigen-Ab2-coated magnetic-responsive micro-bead complexes52 (FIG. 8B). The formation of Ab1-antigen-Ab2-coatedmagnetic-responsive micro-bead complexes 52 anchors the boundantibody-coated magnetic-responsive micro-beads 50 onto the surface ofthe reaction chamber 34. After all antigen molecules 48 have reacted toform the Ab1-antigen-Ab2-coated magnetic-responsive micro-bead complexes52, any unbound, i.e., free, antibody-coated magnetic-responsivemicro-beads 50 exit the reaction chamber 34 via a chamber outlet 38(FIG. 8C), leaving behind the bound antibody-coated magnetic-responsivemicro-beads 50 in the reaction chamber 34. FIG. 9 shows the formation ofthe Ab1-antigen-Ab2-coated magnetic-responsive micro-bead complex 52 asa liquid sample containing an antigen 48 flows through the reactionchamber 34 and rehydrates the antibody-coated magnetic-responsivemicro-beads 50.

A negative liquid sample, i.e., a sample not containing detectabletraces of the targeted analyte, results in zero antibody-coatedmagnetic-responsive micro-beads anchored to the reaction chamber'ssurface, as the Ab1-antigen-Ab2-coated magnetic-responsive micro-beadcomplexes cannot form. An analyte (i.e., antigen)-positive sample, onthe other hand, results in antibody-coated magnetic-responsivemicro-beads anchored to the reaction chamber's surface via theAb1-antigen-Ab2-coated magnetic micro-bead complexes. Thus, the higherthe concentration of analyte in the liquid sample, the greater thenumber of magnetic-responsive micro-beads anchored to the reactionchamber's surface, and hence the fewer the number of freemagnetic-responsive micro-beads reaching the reaction chamber assayoutlet. In the extreme case of very high analyte concentration, allantibody-coated magnetic-responsive micro-beads will be anchored to thereaction chamber's surface, and none will exit through the reactionchamber's assay outlet.

After flowing through the reaction chamber, the liquid sample,self-propelled by capillary action, reaches the pScreen™ magneticmicro-bead concentration counter portion of the device (which principleof operation has been described previously). If the liquid sampleflowing into the test micro-channel and the calibration micro-channelcontains no magnetic-responsive micro-beads, the flow rate in both thetest and calibration micro-channels will be identical, and thus thesample volume collected in each of the graduated columns will beidentical. The user easily is able to observe that the volume of samplein each of the graduated columns is of equal length. On the other hand,if the sample coming from the micro-channel splitter containsmagnetic-responsive micro-beads in any concentration other than zero,the flow of the liquid in the test micro-channel will be retarded (dueto the magnetically-induced flocculation of the magnetic micro-beads).Hence, the length of the volume of liquid in the test graduated columnwill be less than the length of the volume of liquid in the calibrationgraduated column by an amount proportional to the magnetic-responsivemicro-bead concentration in the volume of liquid flowing into thegraduated columns. In other words, the higher the magnetic-responsivemicro-bead concentration in the liquid reaching the test and calibrationmicro-channels, the greater the difference in the lengths of the volumeof liquid observed in the two graduated columns. The resultingdifference between the volumes of liquid collected in the two graduatedcolumns is easily visible to the naked eye.

In an embodiment of the pScreen™ microfluidic immunoassay device,described in detail above and shown in FIG. 4, the test micro-channel(Cm) is split into three test micro-channels 22, 22′ and 22″ which runparallel to each other. In another embodiment of the pScreen™microfluidic immunoassay device, described in detail above and shown inFIG. 5, the test micro-channel (Cm) is split into four micro-channels 22which run parallel to each other.

FIG. 10 shows the pScreen™ microfluidic immunoassay device 10, accordingto the embodiments of the invention, in which the calibration column isonly partially visible. In this embodiment, the read-out is taken whenthe portion of the calibration column that is visible changes color,i.e., fills up with liquid.

In another embodiment of the present invention, there is provided amethod for detecting and quantifying concentration of an analyte in aliquid sample using the pScreen™ immunoassay device having a pluralityof reaction chambers and a system of micro-channels. The methodcomprises adding a liquid sample to a liquid sample inlet which is incontinuous fluid connection with a reaction chamber inlet manifoldmicro-channel, which is in continuous fluid connection with a pluralityof reaction chambers. Each of the reaction chambers has adsorbed on itssurface a plurality of immobilized antigen-specific antibodies (Ab1)specific to an analyte. The surface of each of the reaction chambersalso has a plurality of magnetic-responsive micro-beads desiccatedthereon, in which each of the plurality of magnetic-responsivemicro-beads is coated with an antigen-specific antibody (Ab2) specificto the analyte. The method further comprises having the liquid sampleincubate within the plurality of reaction chambers, which causesrehydration of the plurality of antibody-coated magnetic-responsivemicro-beads as the liquid sample is added and flows in each of thereaction chambers, which rehydration disperses the antibody-coatedmagnetic-responsive micro-beads in the liquid sample, binding therehydrated antibody-coated magnetic-responsive micro-beads as well asthe antigen-specific antibodies immobilized on the surface of each ofthe reaction chambers to any analyte present in the liquid sample toform Ab1-analyte-Ab2-coated magnetic-responsive micro-bead complexes onthe surface of each of the reaction chambers, and having the liquidsample containing any unbound antibody-coated magnetic-responsivemicro-beads exit the plurality of reaction chambers through each of thereaction chamber's outlet manifold micro-channels. Each of the outletmanifold micro-channels is comprised of a micro-channel that is incontinuous fluid connection with a connector micro-channel, which is incontinuous fluid connection with a terminal flow splitter micro-channel.The terminal flow splitter micro-channel bifurcates to form acalibration micro-channel (Co) and at least one test micro-channel (Cm).Each of the outlet manifolds is connected to the connector micro-channelvia a passive valve. Each of the passive valves function tosubstantially stop, for a given period of time, the flow of fluid frommoving forward into the micro-channel splitter, hence allowing time forany antigen, capture antibodies and magnetic-responsive micro-beads tochemically react in the plurality of reaction chambers. The period oftime for which each of the passive valves substantially stops the fluidfrom moving forward is determined by a desired incubation time, whichincubation time provides an optimum reaction time for the specificantibody-antigen of the assay.

The method further comprises measuring the volume, Vm, of the liquidsample passing through the at least one test micro-channel exposed to amagnetic field gradient, in which the presence of any unboundantibody-coated magnetic-responsive micro-beads in the at least one testmicro-channel which is exposed to the magnetic field gradient causesflocculation of the antibody-coated magnetic-responsive micro-beads inthe liquid sample which reduces the flow rate of the liquid samplethrough the at least one test micro-channel, and the volume of theliquid sample passing through the calibration micro-channel, Vo, notexposed to a magnetic field gradient, calculating the difference Vo−Vm,or the ratio (Vo−Vm)^(p)/(Vm)^(q), wherein p and q are derived through acalibration process, wherein the ratios Vm/Vo and (Vo−Vm)^(p)/(Vm)^(q)are a proxy for the number of magnetic-responsive micro-beads in theliquid sample exiting the reaction chambers, which is a proxy for theconcentration of analyte in the liquid sample, and then quantifying theconcentration of analyte in the liquid sample.

In another embodiment of the present invention, there is provided asingle use, portable, lab-on-card microfluidic pScreen™ immunoassaydevice having a plurality of reaction chambers and a system ofmicro-channels for detecting and measuring an analyte in a liquidsample.

In particular, as shown in FIG. 16A (top view) and 16B (side view), themicrofluidic pScreen™ immunoassay device 10 comprises the microfluidicdevice 10 with a sealing layer 72 atop the microfluidic device (shown inFIG. 16B). As shown in FIG. 16A, a liquid sample inlet 8 is provided foraccepting a liquid sample. The liquid sample inlet 8 is in continuousfluid connection with a reaction chamber inlet manifold micro-channel 89which is in continuous fluid connection with a plurality of reactionchambers 34. Each of the reaction chambers 34 has adsorbed on itssurface a plurality of immobilized antigen-specific antibodies (Ab1)specific to an analyte, as well as having a plurality ofmagnetic-responsive micro-beads desiccated thereon. In an embodiment,each of the plurality of reaction chambers 34 may be coated with adifferent Ab1 antibody as well as having different Ab1 surface densitiesin order to react with a plurality of antigens. Each of the plurality ofmagnetic-responsive micro-beads is coated with an antigen-specificantibody (Ab2) specific to the analyte.

In an embodiment, magnetic-responsive micro-beads may be coated withdifferent Ab2 antibodies in order to react with a plurality of antigens.Flow of the liquid sample through each of the plurality of reactionchambers 34 rehydrates the plurality of antibody-coatedmagnetic-responsive micro-beads, which disperses into the liquid sample.The analyte present in the liquid sample binds to the dispersedantibody-coated magnetic-responsive micro-beads as well as to theantigen-specific antibodies immobilized on the surface of each of theplurality of reaction chambers 34 to form Ab1-analyte-Ab2-coatedmagnetic-responsive micro-bead complexes. Any unbound antibody-coatedmagnetic-responsive micro-beads, dispersed in the fluid, exit thereaction chambers 34 through outlet manifold micro-channels 68, 68′ 68″,flow through a connector micro-channel 67, and then flow through aterminal flow splitter micro-channel splitter 19 that bifurcates to forma calibration micro-channel (Co) 20 and at least one test micro-channel(Cm) 22. The calibration micro-channel 20 and the at least one testmicro-channel 22 each are kept at an equal and constant pressure.

Each of the plurality of the reaction chamber's 34 outlet manifoldmicro-channels 68, 68′, 68″ is in continuous fluid connection with theconnector micro-channel 67 via a passive valve 66, 66′, 66″. As usedherein, the term “passive valve” is meant to describe a valve that doesnot have any moving mechanical parts. The passive valves 66, 66′, 66″are formed by having sharp edges at the intersections of the pluralityof outlet manifold micro-channels 68, 68′, and 68″ with the connectormicro-channel 67, as shown in FIG. 17A and 17B (top views) and 17C(perspective view). Fluid flow substantially stops to a negligible flowrate at each intersection of a passive valve 66, 66′, 66″ with theconnector micro-channel 67. The fluid resumes flowing into the connectormicro-channel 67 only when the valves 66, 66′, 66″ burst sequentially,each within a few seconds from one other. The term “burst” is awell-known term in the field of microfluidics, and is described indetail below. The period of time that the fluid is substantially stoppedat the passive valves 66, 66′, 66″ is referred to as “the delay time,”because the valves delay the time it takes for the fluid to move fromthe outlet manifold micro-channels 68, 68′, 68″ into the connectormicro-channel 67.

FIG. 17C shows the spatial relationship of one outlet manifoldmicro-channel 68, a passive valve 66 located at the end of the outletmanifold micro-channel 68, and the connector micro-channel 67. Thebottom surface 84 of the outlet manifold micro-channel 68 forms a90-degree angle with the side wall 85 of the connector micro-channel 67.The side wall 86 of the outlet manifold micro-channel 68 forms a90-degree or greater angle 87 with the side wall 85 of the connectormicro-channel 67. These 90-degree or greater angles have the effect ofincreasing the contact angles of the fluid against the walls of themicro-channels, above 90 degrees, which creates a convex meniscus,which, as per the Young-Laplace principle, creates a pressure gradientin the direction opposite to the fluid direction of motion. This, inturn, stops the flow of fluid from proceeding forward through theintersection of the outlet manifold micro-channel and the connectormicro-channel, hence creating a passive valve to the fluid forwardmotion. Once the flow is stopped, it requires an applied pressuregradient, in the direction of the fluid motion, which is greater thanthe pressure gradient created by the passive valves, to get the fluid toresume flowing past the valve and into the connector micro-channel 67.This process would normally require a mechanical action to apply thepressure gradient in the direction of the flow, either automatically byusing an electrical or mechanical timer system, or by anoperator-activated system.

In this embodiment, a primary flow splitter 74 (shown in FIG. 16) isplaced before valve 66 joins the connector micro-channel 67. The primaryflow splitter forms a secondary micro-channel, referred to as a delaymicro-channel 65, which connects to the connector micro-channel 67 andis located before valve 66 intersects with the connector micro-channel67 along the flow direction. As used herein, the area where the valves66, 66′, 66″ intersects with the connector micro-channel is referred toas “the valve-connector intersection”.

As described above, fluid flow is stopped where the valve 66 intersectswith the connector micro-channel 67, whereas the fluid in the delaymicro-channel 65 continues to flow via capillary action. Fluid from thedelay micro-channel 65 therefore is free to enter the connectormicro-channel 67, as the delay micro-channel 65 forms a continuoussurface with the connector micro-channel's 67 wall. This is in contrastto the intersection between the micro-channel outlet manifold 68 and theconnector micro-channel 67, which forms the passive valve 66 due totheir sharp edges. As used herein, a sharp edge is defined as the angleformed between the tangents of two surfaces, at the surfaces'intersection area or line, which angle is greater than zero. As fluidenters the connector micro-channel 67, the fluid wets the walls of theconnector micro-channel 67 and continues to flow by capillary action. Asfluid reaches the first valve 66 in its path, it joins the fluideffectively stopped at the intersection of the valve 66 and theconnector micro-channel 67, providing a continuous surface for the fluidto move past the valve 66. At this point, the valve 66 is said to burst,since the fluid is now flowing through the valve propelled by capillaryaction of the fluid meniscus that travels through the connectormicro-channel 67.

The period of time that it takes for the fluid to travel through thedelay micro-channel 65 is referred to herein as the delay time. Duringthe delay time, the flow in each of the plurality of reaction chambers34, except for the reaction chamber connected to the delaymicro-channel, is stopped, whereas the flow in the reaction chamberconnected to the delay micro-channel can be reduced to an arbitrary lowlevel, or substantially stopped, i.e., the flow rate is decreased to anegligible rate, as described in detail below. This delay time allowstime for the magnetic-responsive micro-beads to bind more effectively tothe captured antigen. The delay time is chosen to match a desiredincubation time value, which value is determined as providing theoptimum reaction time for the specific antibody-antigen of the assay. Asused herein, the optimum reaction time is defined as the shortest timethat results in between about 60% to 90% of the total number of antigenmolecules in the fluid contained in the plurality of reaction chambersto bind to the capture antibodies.

As shown in FIG. 16A, the calibration micro-channel 20 is in continuousfluid connection with one or more parallel calibration graduated columns26, and the at least one test micro-channel 22 is in continuous fluidconnection with one or more parallel test graduated columns 26′. The atleast one test micro-channel 22 is exposed to a magnetic field gradient,which causes flocculation of the magnetic-responsive micro-beads in theat least one test micro-channel 22. The flocculation reduces the flowrate ({tilde under (O)}m) of the liquid sample in the at least one testmicro-channel 22 compared to the flow rate ({tilde under (O)}o) of theliquid sample in the calibration micro-channel 20. Each of the one ormore test graduated columns 26′ has a graduated scale 30 thereon whichprovides a read-out of the sample volume collected in each of the testgraduated columns 26′, Vm, and the calibration graduated columns 26, Vo.The difference Vo−Vm, or the ratio (Vo−Vm)^(p)/(Vm)^(q), wherein p and qare derived through a calibration process, are a proxy for the number ofmagnetic-responsive micro-beads in the liquid sample, which is a proxyfor the concentration of analyte in the liquid sample, which allows forthe quantification of the concentration of analyte in the liquid sample.

In an embodiment, the reaction chambers 34 are about 20 mm to 50 mmlong, about 1 mm to 5 mm wide, and about 50 μm to 200 μm deep. The inletmanifold micro-channel 89 is about 50 μm to 300 μm deep and wide, andabout 10 mm to 40 mm long. The outlet manifold micro-channels 68, 68′,and 68″ are about 50 μm to 200 μm wide, and about 50 μm to 300 μm deep.The connector micro-channel 67 is about 50 μm to 400 μm wide, and about60 μm to 400 μm deep. The delay micro-channel 65 about 30 mm to 120 mmlong, about 50 μm to 300 μm wide, and about 50 μm to 300 μm deep.

The delay time can range between about 30 seconds and 10 minutes. In anembodiment, the delay time ranges between about 1 minute and 4 minutes.

In an embodiment, the number of reaction chambers is between 1 and 8. Inanother embodiment, the number of reaction chambers is 4.

In another embodiment, as shown in FIG. 18, each of the reactionchambers 34, except for the reaction chamber 34 having the primary flowsplitter 74 and the delay micro-channel 65, are in continuous fluidconnection with secondary flow splitter micro-channels 75, which are incontinuous fluid connection with the outlet manifold micro-channels 68′,68″. Each secondary flow splitter 75 is in continuous fluid connectionwith an appendix micro-channel 69, which terminates in a vent port 70that is open to atmospheric pressure. The number of appendixmicro-channels therefore is equal to the number of reaction chambersminus one. The appendix micro-channels 69 are about 50 μm to 400 μmwide, about 50 μm to 400 μm deep, and about 30 mm to 120 mm long.

Whereas fluid flow stops where the outlet manifold micro-channels 68,68′, 68″ intersect the connector micro-channel 67, fluid flow iscontinuous in the appendix micro-channels 69. The advantage ofintroducing the appendix micro-channel in the present invention isevident when a hydrophilic sealing layer is used in place of ahydrophobic sealing layer, which sealing layer is described in detailbelow.

The passive valves have three sharp edges and one continuous surface,which is a sealing layer. FIG. 19 shows the sealing layer 72, which canbe made of either a hydrophilic or hydrophobic material. The advantageof a hydrophilic sealing layer is that it provides, in addition to thethree walls of the micro-channels which are carved in the microfluidicdevice, a fourth continuous surface which promotes fluid flow movementby capillary action throughout the device, thus increasing the overallflow rate and reducing the assay time. In contrast, a hydrophobic layerresists the flow, thus reducing the capillary action of the three wallsof the micro-channels carved in the microfluidic device 10. The term“carved” is used herein to refer to any of the methods used to make themicrofluidic device of the present invention, as described below. Usinga hydrophilic material may, however, result in failure of the passivevalves to stop the fluid flow through the valves. This is because asealing layer made of a hydrophilic material, by providing a hydrophiliccontinuous surface, allows the fluid to flow easily past the passivevalves, thus causing the valves to fail, i.e., the fluid is not stoppedat the valves. This is in contrast to a hydrophobic sealing layer, whichimpedes the forward flow of fluid, and thus acts as a flow stop inaddition to the three walls of the micro-channels. Unimpeded flow occursanytime the hydrophilic sealing layer produces a capillary pressurelarger than the pressure gradient created by the sharp edges of thepassive valves which opposes the flow through the valves. Thus, with ahighly hydrophilic surface layer this unimpeded flow will very likelyoccur. However, incorporating the appendix micro-channels 69 (shown inFIG. 18) in the present invention, a hydrophilic sealing layer can beused without causing failure of the valves 66, 66′, 66″. This is becausewhen fluid is flowing in the appendix micro-channels 69, the pressuregradient across the valves 66, 66′, 66″ is reduced by the fluid flowcapillary action in each of the appendix micro-channels, which creates apressure gradient in the opposite direction that is larger than thepressure gradient created by the capillary action through the valves,hence stopping the flow of the fluid from flowing through the passivevalves.

In accordance with the invention, during the delay time, the flow ineach of the plurality of reaction chambers 34 is reduced to an arbitrarylow level, or is substantially stopped, i.e., the flow rate is decreasedto zero or is negligible. This arbitrary low flow rate is achieved byadjusting the length, width, and depth of the delay micro-channel andthe appendix micro-channels. Because both the delay micro-channel andthe appendix micro-channels are relatively narrow and longmicro-channels, flow resistance in these micro-channels rapidlyincreases, which in turn causes the flow rate (which is inverselyproportional to the fluid resistance) in these micro-channels todecrease and stop. In addition, by increasing the width and depth of thelast few mm of the final portions of the delay micro-channel and theappendix micro-channels (which represents only a small portion of thetotal length of these micro-channels), by about two to six fold, thecapillary pressure in these micro-channels is greatly reduced. Thus,because the flow resistances in both the delay micro-channel and theappendix micro-channels are substantially unchanged, by increasing thewidth and depth of the final portion of these micro-channels, thisresults in a further reduction of the flow rate. Hence, by adjusting thelength, width and depth of the delay micro-channel and the appendixmicro-channels, the flow rate in the reaction chamber can be madearbitrarily low so that flow substantially stops.

Referring again to FIG. 19, the immunoassay device is constructed on twolevels, rather than in a co-planar configuration, with the one or morereaction chambers 34 constructed on a bottom level, the system ofembedded micro-channels 78 constructed on a top level, and a sealinglayer 72 positioned in between the top and bottom levels. In particular,all of the micro-channels of the device are placed on one level by moldcasting or any of the above-described methods. The micro-channels areformed by sealing two layers of materials, one in which the system ofmicro-channels are embedded 78 and the other forming the sealing layer72. The one or more reaction chambers 34 which are formed on the lowerlevel is constructed by bonding the bottom plate 88 of each of thereaction chambers 34 (which are coated with one or more captureantibodies) to the bottom side of the sealing layer 72 to form the oneor more reaction chambers 34. The two layers, i.e., the embeddedmicro-channels 78 and the sealing layer 72, are connected and sealed byadhesives or other bonding materials. Bonding the one or more reactionchambers 34 with the bottom of the sealing layer 72 may be accomplishedby a variety of methods, which include, without limitation, doublepressure sensitive adhesive 71 or thermal bonding. Holes which form thereaction chamber inlet 36 and the reaction chamber outlet 38 (also shownin FIG. 8) are placed on the sealing layer to transfer fluid in and outof the one or more reaction chambers 34.

Referring now to FIG. 20A and B, the liquid sample inlet 8 of theimmunoassay device 10 has a conical shape (shown in cross-section inFIG. 20A). The inner surface 76 of the liquid sample inlet 8 is madesuper-hydrophobic by either coating the inner surface 76 with asuper-hydrophobic layer or by a surface texture that creates a capillarycontact angle 77 between 120 degrees and 160 degrees. The upper opening90 of the liquid sample inlet 8 has a diameter between about 3 mm and 10mm. The lower opening 79 of the liquid sample inlet 8 (which iscontinuous with the reaction chamber inlet; not shown) is between about1 mm and 6 mm. The cone angle is designed to match the capillary contactangle. The inlet shape is designed and the surface is made hydrophobicto create a convex meniscus 82 that creates a pressure within the liquidsample that is greater than atmospheric pressure and, by virtue of theYoung-Laplace's principle, creates a pressure gradient that favors theflow of the liquid sample through the microfluidic device. This furtherincreases the liquid sample flow rate through the device by providing anadditional capillary pressure gradient in the direction of the flow fromthe liquid sample inlet toward the system of micro-channels of thedevice of the present invention.

As shown in FIG. 20B (expanded view of the left side of the inletopening), the cone angle (alpha) 80 of the conical inlet is related tothe fluid contact angle (beta) 77, which is the angle formed by inletsurface tension force, 81, with the inlet hydrophobic inner surface. Fora fixed value of beta, the optimum value of alpha is given by thefollowing equation: alpha=beta−90°. When this equation is satisfied,then the capillary pressure created by the fluid meniscus is maximum,and is equal to two times the surface density divided by the radius ofthe inlet circumference at the location where the fluid meniscusintersects the conical inlet's inner surface. Therefore, as the fluid inthe inlet is reduced as fluid moves into the device, the radius of theinlet circumference, at the location where the fluid meniscus intersectsthe conical inlet's inner surface, decreases, and the capillary pressurecreated by the meniscus increases.

In an embodiment, the walls of the micro-channels are coated withhydrophilic films to decrease the fluid contact angle and in turnincrease the capillary action and thus the flow rate, hence reducing theassay time. Hydrophilic films used in coating include, but are notlimited to, hyaluronic acid, bovine serum albumin, or a protein-freeblocking solution. Other hydrophilic liquid or sprays also may be used.

In an embodiment, the micro-beads are dispensed in the one or morereaction chambers in nanoliter-sized drops 83 (shown in FIG. 23). Thenanoliter-sized drops are dispensed using methods which include, withoutlimitation, micro-array spotting pins or a nanoliter-sized pipettemechanism. Drops may be dispensed in a variety of configurations andarrays. Drop size may range from about 5 nl to greater than 100 nl.Drops may be layered directly on top of one another in one or morelayers to increase the number of micro-beads in each spot and overall inthe one or more reaction chambers, which promotes micro-bead releasewhen a liquid sample rehydrates the micro-beads. A higher percentage ofmicro-beads releases from the surface of the one or more reactionchambers when more than one layer of drops is applied, and thus moremicro-beads are rehydrated and become available for binding and fordetection of analyte in the liquid sample.

In an embodiment, sucrose, glycerol, and trehalose or any combination oftwo or more of sucrose, glycerol and trehalose, are added to themicro-bead's buffer solution to promote micro-bead release. Themicro-bead's buffer solution typically is phosphate buffered saline(PBS). When sucrose, glycerol, trehalose, or a combination thereof isadded to the micro-bead buffer solution, a higher percentage ofmicro-beads releases from the surface than when sucrose, glycerol,trehalose are not included in the solution. Therefore, when sucrose,glycerol, trehalose, or a combination thereof is added to the micro-beadbuffer solution, more micro-beads are rehydrated and thus available forbinding and for detection of analyte in the liquid sample. Concentrationof these additives from about 5-40% of the buffer solution promotesmicro-bead release.

In an embodiment, the dispensing solution density is altered dependingon the density of micro-beads used. Micro-beads have a higher densitythan both water and PBS. Thus, the micro-beads settle to the bottom ofthe reservoir or tubing during dispensing and the number of micro-beadsdispensed per drop changes over time. In order to keep the number ofmicro-beads dispensed per drop constant over time, the density of themicro-beads must be the same as the density of the solution in which themicro-beads are suspended during dispensing to achieve a neutralbuoyancy micro-bead dispensing solution. Density of the solution can bealtered to match the density of the micro-beads using additives thatincrease solution density, which include but are not limited to heavywater, glycerol, sucrose, polyethylene glycol, or a combination of twoor more of these additives. Concentrations of these additives can bevaried to match the density of a given micro-bead. In particular, theuse of heavy water is very advantageous because it simply replaces thewater in the micro-bead dispensing solution. Suspension in neutralbuoyancy fluid prevents settling and maintains uniform micro-beadconcentration in the dispensing buffer for the entire time it takes todispense the drops. Dispensing can take several hours depending on thenumber of devices prepared. This allows for uniformity of micro-beadconcentration within each reaction chamber and between reactionchambers.

In the above-described embodiments of the method and device of thepresent invention, a liquid sample flows via capillary action thoroughthe plurality of the reaction chambers and the system of micro-channels.The flow rate of the liquid sample varies with the position of the fluidmeniscus in accordance with the embodiments of the invention. As theliquid sample enters each reaction chamber, the flow rate is relativelyhigh, typically between about 0.3 μl/sec to 0.4 μl/sec, then the flowrate decreases as the fluid moves forward via capillary pressure createdby the fluid meniscus. The mean flow rate in the reaction chamber isbetween about 0.15 μl/sec to 0.2 μl/sec. As the fluid meniscus entersthe outlet manifold micro-channels, the flow rate further decreases,according to well-known microfluidic principles, which, in brief, statesthat the pressure gradient created by a micro-channel is inverselyproportional to the size of the micro-channel. And, because the depth ofeach outlet manifold micro-channel is larger than the depth of eachreaction chamber, the meniscus in each of the outlet manifoldmicro-channels creates a lower pressure gradient and hence a lower flowrate. Then, as the liquid splits at the primary and secondary flowsplitter micro-channels, the meniscuses of the branches connected to thepassive valves stop, while the meniscus of the fluid in the delay andappendix micro-channels continue to move forward. Because both the delaymicro-channel and the appendix micro-channels are narrow and longmicro-channels, their flow resistance rapidly increases, which in turncauses the flow rate (which is inversely proportional to the fluidresistance) to rapidly decrease and substantially stop. Then, once themeniscus enters the connector micro-channel causing the system ofpassive valves to burst, the flow bypasses the appendix micro-channelsand the delay micro-channel, and moves through the valves, and sinceboth appendix micro-channels and the delay line have a higher flowresistance than each the outlet manifold micro-channels, the flow rateincreases. Then, as the fluid flows through the calibrationmicro-channel (Co) and at least one test micro-channel (Cm), the flowrate slightly decreases since these micro-channels introduce a high flowresistance. Finally, the fluid enters the graduated columns, and sincethe graduated columns' width and depth are larger than the test andcalibration micro-channels, the capillary pressure at the fluid meniscusis lowered and the flow rate decreases. When the fluid is in thegraduated columns, the flow rate remains substantially constant, sincethe flow resistance of the graduated columns is much smaller than thesum of the flow resistances of all the preceding micro-channels. Theinitial high flow rate found in the reaction chambers favors there-hydration of the magnetic micro-beads. The substantially stopped flowrate, when the fluid is in the appendix micro-channels and in the delayline, favors antigen incubation and formation of themicro-beads-antigen-capture antibody complex. Then, the increase in flowrate, when the fluid is in the connector micro-channel, favors themovement of the magnetic micro-beads that are not bound to the surfacesof the reaction chambers via the antigen. This step reduces the numberof micro-beads that may bind non-specifically to the surfaces of thereaction chambers, and thus reduces assay background. Finally, thesubstantially constant flow rate encountered when the meniscus is in thegraduated columns provides a uniform flux of micro-beads out of thereaction chambers and through the test and control micro-channels.

In another embodiment, there is provided a device for the quantificationof DNA fragments based on flocculation of silica micro-beads. Underspecific conditions, DNA fragments bind to the surface of the silicamicro-beads and trigger their aggregation into bead clusters (i.e.aggregates). These clusters are held together by an electrostaticinteraction between the DNA fragments and the silica micro-bead surface.Accordingly, the degree of flocculation is proportional to the amount ofDNA fragments present in a sample. A higher concentration of DNAfragments will lead to more and larger aggregates, while a lowerconcentration of DNA fragments to smaller and fewer clusters. By mixingmagnetic-responsive silica micro-beads to nonmagnetic-responsive silicamicro-beads, the DNA-induced clusters will become magnetic-responsive.As such, even a small amount of magnetic-responsive silica micro-beadswill lead to large magnetic-responsive bead clusters by addition ofnon-magnetic-responsive silica micro-beads.

To trigger and facilitate the binding of DNA fragments to the silicamicro-beads and subsequent formation of bead clusters, the samplecontaining the DNA fragments first is mixed with a low pH buffer thatpromotes the binding of the DNA strands onto the silica micro-beads byelectrostatic interaction. Subsequently, a chaotropic salt-based buffercontaining either isopropyl or ethanol alcohol is added, which promotesthe formation of the bead clusters by hydrophobic interactions.Following the mixing of the sample with the clustering-promotingsolutions, an aliquot of the sample is transferred into a custom pScreenmicrofluidic device, shown in FIG. 25, which first promotes collisionsbetween silica micro-beads, and hence the formation ofmagnetic-responsive bead clusters. Then, it splits the flow into twocalibrated micro-channels, one of which is placed adjacent to apermanent rare-earth magnet. All magnetic-responsive bead clustersexiting the chamber and entering the micro-channel will aggregate in theproximity of the magnet and create a localized restriction, whichretards the flow of liquid. Conversely, the flow in the other channel isunaffected by the presence of clusters. This enables visual readout ofthe concentration of DNA fragments in the sample by comparing thevolumes of liquid accumulated into the two channels during an assay runtime described above. In the absence of DNA fragments in the sample,only the magnetic-responsive silica micro-beads will aggregate in theproximity of the magnet. Therefore, by keeping the fraction ofmagnetic-responsive silica micro-beads low, a sample absent of DNAfragments will result in no appreciable difference in the amount of thevolumes of liquid accumulated in the two channels. On the other hand, ifDNA fragments are present in the sample, they will trigger the formationof large magnetic-responsive clusters, which then will create a largeaggregate in the proximity of the magnet and introduce a localizedrestriction, which retards the flow of liquid in that column (i.e. testcolumn).

The pScreen microfluidic device, therefore, allows for rapid, simple andinexpensive quantification of DNA concentration in a sample. The devicedoes not require complex instrumentation and can be performed in lessthan three minutes. Moreover, it is compatible with complex samplesincluding, without limitation, unpurified PCR amplification products,and thus can be expected to seamlessly integrate into various commonmolecular biology techniques and workflows.

In one embodiment of the device for the quantification of DNA fragmentsbased on flocculation of silica micro-beads, there is provided a singleuse, portable, lab-on-card microfluidic device (shown in FIGS. 25A,B),comprising a microfluidic card 10 and a sealing layer 7 (shown in FIG.25B) and a liquid sample inlet 8, defined by an opening for acceptingthe liquid sample. The microfluidic device has a reaction chamber 34 anda system of secondary micro-channels for detection and quantification ofDNA fragments on flocculation of silica micro-beads in themicro-channels and the detection of magnetic aggregates in a liquidsample. As shown in FIG. 25A, the liquid sample inlet 8 is in continuousfluid connection a reaction chamber 34, which is in continuous fluidconnection with a micro-channel splitter 18, which bifurcates to form acalibration micro-channel (Co) 20 and one test micro-channel (Cm) 22.The calibration micro-channel 20 is in continuous fluid connection witha control graduated column 26′, and the test micro-channel 22 is incontinuous fluid connection with a graduated parallel test graduatedcolumn 26′. Both columns are vented by small venting holes, 100 and 100′(shown in FIG. 25A, B), placed at the end of each column. The testmicro-channel 22 is exposed to a magnetic field gradient, which causesflocculation of the magnetic-responsive clusters of the silicamicro-beads in the test micro-channel 22, in the proximity of themagnet.

In use, the microfluidic device includes an assay solution comprised oftwo liquid buffers, Buffer 1 containing Reagent 1, and Buffer 2containing Reagent 2. Buffer 1 contains a mixture of magnetic- andnonmagnetic-responsive silica micro-beads. First, an aliquot of the DNAsample to be tested is mixed with Buffer 1 and incubated for a period oftime, T1. Then, Buffer 2 is added to the DNA sample and Buffer 1, mixed,and then incubated for a period of time, T2, to promote the formation ofmagnetic-responsive bead clusters. This mixture, comprised of the DNAsample, Buffer 1 and 2, then is placed on the inlet 8 of themicrofluidic pScreen™ device 10. The magnetic-responsive bead clusters,dispersed in the liquid sample, flow through the reaction chamber 34 andexit through the micro-channel splitter 18 that bifurcates to form thecalibration micro-channel (Co) 20 and the test micro-channel (Cm) 22.The flocculation of the magnetic-responsive clusters reduces the flowrate ({tilde under (O)}m) of the liquid sample in the test micro-channel22 compared to the flow rate ({tilde under (O)}o) of the liquid samplein the calibration micro-channel 20. The test graduated column 26 has agraduated scale 30 thereon which provides a read-out of the samplevolume collected in the test graduated column 26, Vm and the controlgraduated column 26, Vo, in which the difference, Vo−Vm, or thedifference, Lo−Lm, where Lo is the length of the control column, and Lmthe length of the test column, or the ratio (Vo−Vm)p/(Vm)q, wherein pand q are derived through a calibration process, is a proxy for thenumber of magnetic-responsive bead clusters in the liquid sample, whichis a proxy for the concentration of DNA fragments in the liquid sample.

In an embodiment, the unknown concentration of DNA fragments in a sampleis derived by comparing the difference, Lo−Lm, to the calibration curveobtained for a DNA fragment of about equal length, i.e., with an equalnumber of base pairs.

In an embodiment, Buffer 1 is a low pH water-based buffer containingbetween 1 Mole/Liter and 2 Mole/Liter of acetic acid (CH₃COOH), 2Mole/Liter and 3 Mole/Liter of sodium chloride (NaCl), and between 5%v/v and 40% v/v polyethylene glycol, PEG 400; and contained silicanonmagnetic-responsive micro-beads in concentrations between 10 μg/μland 60 μg/μL, and magnetic-responsive silica beads in concentrationsbetween 1.0 μg/μl and 6.0 μg/μl. Buffer 2 is a salt and alcohol-basedbuffer containing isopropyl alcohol between 10% and 50% v/v, guanidinehydrochloride between 10% and 30% w/v, and sodium perchlorate between 2%and 10% w/v. Alternatively, ethanol between 20% and 75% v/v can be usedin place of isopropyl.

In an embodiment, Buffer 2 contains dyes in concentration between 0.2and 0.5 mg/ml to make the otherwise clear liquid readily visible to thenaked eye. Dyes may include Trypan Blue, Rhodamine 6G, and CrystalViolet.

In an embodiment, the non-magnetic-responsive silica micro-beads inBuffer 1 vary in size between 1 μm and 12 μm, and themagnetic-responsive micro-beads in Buffer 1 contain a magnetic kerneland are coated with a silicon dioxide (SiO₂) layer and vary in sizebetween 2.5 μm and 4.5 μm.

In an embodiment, the reaction chamber 34 is 10 mm to 50 mm long, 1 mmto 5 mm wide, and 50 μm to 200 μm deep; the inlet and outlet manifolds 8are 50 μm to 300 μm deep and wide and 10 mm to 40 mm long; themicro-channels 20 and 22 have a width between 50 μm and 200 μm, and adepth between 50 μm and 300 μm; and the columns 26 and 26′ have a widthbetween 0.5 mm and 2 mm, a depth between 50 μm and 400 μm, and a lengthbetween 100 mm and 500 mm. The sealing layer can be made of either ahydrophilic or hydrophobic material. The advantage of using ahydrophilic, rather than a hydrophobic material, is that a hydrophilicsealing layer provides, in addition to the three walls of allmicro-channels embedded in 10, a fourth continuous surface that promotethe flow motion by capillary action throughout the device, henceincreasing the overall flow rate and reducing the assay time.

Passive valves are essential features of many microfluidic devices. Theyenable the—on/off switch of flow without the assistance of any externaldevices, battery or mechanical transducers. For example, in theembodiment shown in FIG. 25, the user is required to either measure atthe same time, both Vm and Vo, or alternatively measure either Vm or Vowhile measuring the total volume added into the device inlet.Furthermore, too large a sample volume may fill both columns to fullcapacity. As the control column reaches its end, the flow will continuein the test channel, eventually reaching the same length of the controlchannel, hence resulting in a test failure. On the other hand, a smallsample volume will result in both Vm and Vo being shorter than optimum,resulting in a smaller difference between test and control columns,which in turn will reduce assay sensitivity. Here, we present a novelmicrofluidic passive valve that is able to stop the flow in a primarymicro-channel as the liquid flowing in a secondary micro-channel reachesa predetermined mark.

In another embodiment, shown in FIG. 26, the columns 26 and 26′ arejoined together by a joining channel, 101, where a small venting hole102, acting as a passive valve, is placed through the sealing layer.This passive valve provides access to atmospheric pressure to the fluidinside both columns, and thus allows the flow to move independently inboth columns under the capillary pressure created by the flow meniscusof the liquid in both columns. The location of the passive valve hole isplaced so that the volume of the control column from the entrance of thecontrol column 103′ to the venting hole location 102 is larger than thevolume of the test column from the entrance of the test column 103 tothe venting hole location 102. Thus, the liquid flowing in the controlcolumn will reach the venting hole first, as the liquid in the testcolumn can only flow at the same speed or slower than the liquid in thecontrol column. As soon as the liquid in the control column reaches theventing hole 102, the liquid itself seals the venting hole, creating anair pocket between the flow from the control column and the flow frontof the test column, and thus stopping the liquid in the test column frommoving forward. The liquid from the control column is prevented fromexiting the venting hole, as the top surface of the sealing layer ishydrophobic or partially hydrophobic. Thus, in this embodiment, thecontrol column always fills up only to the passive valve 102, at whichpoint the test column stops. Therefore, since the volume Vo is aconstant, Vm is a proxy for the number of magnetic-responsive beadclusters in the liquid sample, which is a proxy for the concentration ofDNA fragments in the liquid sample. The clear advantage is that in thisembodiment only Vm needs to be recorded, and the total sample volumedoes not need to be measured, since the passive valve 102 ensures thatthe assay will always end regardless the volume of the sample added, aslong as the volume exceeds the minimal requirement.

In an embodiment, the venting hole has a diameter that ranges between0.2 mm to 1.0 mm.

In an embodiment, the top surface 72 of the sealing layer around theventing hole 102 has a contact angle of between 40° and 80°.

In an embodiment, the top surface 72 of the sealing layer around theventing hole 102 is treated with a hydrophobic coating to increase thecontact angle above 90°, hence further preventing liquid with surfacetension lower than water from exiting the passive valve 102 by capillaryaction.

The devices in accordance with the embodiments of the invention, may befabricated by methods which include, without limitation, etching each ofthe micro-channels on a plastic substrate using a laser etcher system,by injection mold casting in plastic, and then sealing the top of eachof the micro-channels with a plastic layer by thermal bonding, or byemploying hydrophilic pressure-sensitive tapes. Suitable plasticsubstrates, plastic sealing and injection mold casting plastics include,without limitation, poly(ethylene terephthalate) glycol,poly(lactic-co-glycolic acid) and poly(methyl methacrylate).

EXAMPLES

The present invention is more particularly described in the followingnon-limiting examples, which are intended to be illustrative only, asnumerous modifications and variations therein will be apparent to thoseskilled in the art.

Example 1—Scientific Basis and Technology Feasibility of the PresentInvention (A) Introduction

The data presented herein describe the effect that flocculation ofmagnetic-responsive micro-beads in a micro-channel has on the flowresistance of liquid in the micro-channel. A liquid seeded withmagnetic-responsive micro-beads in a micro-channel that is exposed to amagnetic field gradient produces a localized micro-bead flocculation.This localized micro-bead flocculation results in a localized reductionof the cross section of the micro-channel, and thus in a localizedincrease of the flow resistance across the flocculation region. Theincrease in resistance, in turn, results in an increased pressure dropacross the flocculation region due to the energy loss in maintaining theflow across the reduced cross-section of the micro-channel. If theexternal forces responsible for the formation of the micro-beadflocculation are stronger than the flow shear-stress on the micro-beadsand their aggregates, the micro-beads' flocculation increases inmagnitude as more incoming micro-beads are added. In the case of aconstant-pressure driven flow (relevant to the present invention), theincreased pressure drop results in a reduction of the micro-channel flowrate. This study investigated and analyzed this phenomenon, and theresults are reported below. These experimental results provide thescientific basis upon which the pScreen™ technology and the presentinvention have been developed.

(B) Experimental Methodology

Experimental data with respect to the effect of magnetic micro-beadflocculation on flow rate in micro-channels are presented. FIG. 1 is anillustration of a constant pressure flow system comprised of twomicro-channels. Test micro-channel 22 (Cm) was exposed to ahigh-magnetic field gradient, while calibration micro-channel 20 (Co)served as a control. The pressure difference driving the flow betweenthe micro-channel inlets 14, 14′ and micro-channel outlets 16, 16′ wasequal between the two micro-channels 20, 22. A liquid sample 12 seededwith a known concentration of magnetic-responsive micro-beads 15 wasadded to both micro-channels 20, 22 and the flow rate in bothmicro-channels 20, 22 was recorded over time. Flocculation of themicro-beads 15 was created by means of a localized high-gradientmagnetic field generated by two magnets 24, 24′.

Experiments were conducted using micro-channels fabricated from glasscapillary tubes having an inner diameter of 50 μm and 100 μm. The lengthof the capillary tubes was varied between 3.0 cm and 7.5 cm. Themagnetic field was generated by two neodymium-iron-boron (NdFeB)permanent magnets 24, 24′ (25 mm×6 mm×1.5 mm; maximum surface field: 0.3T). The magnets 24, 24′ were aligned length-wise along opposite poles.The test micro-channel 22 capillary tube exposed to the magnetic fieldwas placed between the gaps formed between the opposite poles of theNdFeB magnets 24, 24′.

Both micro-channel capillary tubes 20, 22 were partially inserted into arubber stopper portion of a glass vacutainer tube (not shown), leavingabout 0.5 cm of the ends of the capillary tubes visible. Each capillarytube inlet and outlet was inserted in a polystyrene tubing (not shown)having an inner diameter of 360 μm, which tightly fit the 360 μm outerdiameter of the two micro-channel capillary tubes 20, 22. One end of thetubing lead directly to a reservoir 9 containing the liquid sample 12with the magnetic micro-bead solution, and the other end of the tubinglead to the vacutainer tubes which collected the fluid exiting themicro-channel outlets 16, 16′. The sample reservoir 9 was open to theair, and thus was at atmosphere pressure. The vacutainer tubes weresealed and kept under a constant vacuum. The pressure difference betweenthe sample reservoir 9 and the vacutainer tubes induced the liquidsample 12 to flow from the reservoir 9 into the vacutainer tubes. Thepressure difference was maintained at 0.6 mmHg per cm of capillary tubeto provide equal flow rate across capillary tubes of different lengths.Experiments were run in tandem, using two capillary tubes: one for thecalibration, i.e., control, micro-channel 20; and one for the testmicro-channel 22 exposed to the magnetic field gradient. Bothmicro-channel capillary tubes 20, 22 were kept at the same differentialpressure and drew fluid 12 from the same sample reservoir 9. In thecalibration micro-channel capillary tube 20, the sample 12 flowedfreely. In the test micro-channel capillary tube 22, the appliedmagnetic field gradient induced micro-bead 15 flocculation. Thecalibration and test micro-channels 20, 22 were run simultaneously toeliminate common error, such as variation in atmospheric pressure,changes in viscosity due to fluctuation in temperature, and variationsin micro-bead concentration. The suspension medium was 35% (by wt.)glycerol and 65% water to achieve a viscosity similar to that of blood(about 3.6 cP). Green fluorescent dye was added to the suspension mediumto increase visibility of the solution exiting the two micro-channelcapillary tubes 20, 22. Also added to the medium were smooth carboxylmagnetic micro-beads 15 (Spherotech, Inc.) with a diameter of 4.7 μm or8.3 μm. Tests were conducted with a micro-bead 15 concentration between100 micro-beads/μl and 50×10³ micro-beads/μl. Sample volumes werebetween 50 μl to 200 μl and initial flow rates were 0.01 μl/sec. As thesample fluid 12 exited the micro-channel capillary tubes 20, 22, itformed small drops before falling into the vacutainer. The measurementof flow rate was calculated by dividing the drop volume with the timeinterval between drops. The falling drop rate was recorded with a DVDvideo camera. Post video analysis provided the flow rate vs. time.Additional experiments were conducted without a vacutainer. Themicro-channel outlets 16, 16′ were connected to long polystyrene tubing(not shown) which was placed near a graduated ruler. The flow rate wasmeasured by recording the advancement of the fluid meniscus inside thetubing as a function of time. Flow rate values in the glass calibrationmicro-channel capillary tube 20 not exposed to the magnetic fieldgradient were compared with theoretical Hagan—Poiseuille flow {tildeunder (O)}=(πα⁴)ΔP/8 μL) (wherein a is the tube radius, ΔP is thepressure difference across the micro-channel, μ is the fluid viscosity,and L is the tube length) for a fully developed laminar flow of aNewtonian fluid in a cylindrical tube. Additional experiments wereconducted using a micro-channel configuration as shown in FIG. 2 andfabricated in poly(lactic-co-glycolic acid) (PLGA) and Polyethyleneterephthalate by laser etching as above described.

(C) Macroscopic Experimental Results and Data Analysis

FIG. 11 shows the normalized flow rate, i.e., the ratio between the flowrate in the test micro-channel capillary tubes (exposed to the magneticfield gradient) and the calibration micro-channel capillary tubes (notexposed to the magnetic field gradient) versus the number ofmagnetic-responsive micro-beads in the flocculation region. The numberof magnetic-responsive micro-beads is given by the product of the flowrate times the magnetic-responsive micro-bead concentration in thesample. The data show that the normalized flow rate is a monotonicfunction of the number (over three orders of magnitude) ofmagnetic-responsive micro-beads in the flocculation region. The amountby which the flow rate is reduced due to the pressure drop caused by theflocculation depends on the overall capillary length and the size of theflocculation zone. Thus, different aspect ratios of magnetic fieldlength along the micro-channels versus the total length of themicro-channels also were investigated. FIG. 13 shows three data curvesfor different aspect ratios of magnetic field (concentrator) lengthversus the total length of the micro-channels [ratios range from 0.17(triangle) to 0.24 (diamond) to 0.4 (square)]. To the investigators'knowledge, these data are the first to provide direct measurements ofthe effect of magnetic-responsive micro-bead flocculation on fluid flowrate in micro-channels.

(D) Data Analysis

These experimental data demonstrate that over a wide range thenormalized flow rate, {tilde under (O)}m/{tilde under (O)}o, is amonotonic function of the number of magnetic-responsive micro-beadsentering the capillary tubes. What is presented herein is aphenomenological model based on the Poiseuille equation that theinvestigators derived to corroborate these results. The model relatesthe flow rate to the reduction in micro-channel cross-section due to theformation of flocculation. The model predicts the following relationshipbetween flow rate and number of magnetic-responsive micro-beads:

=1/(1+αN),   Equation (9)

where α is a parameter defined below:

α=[(α/R _(eff))⁴−1]/LB,   Equation (10)

B=¾(1−ϵ)(α² −R _(eff) ²)/r ³   Equation (11)

where

is the normalized flow rate, defined as {tilde under (O)}m/{tilde under(O)}o, N is the number of micro-beads in the flocculation, a is thecapillary tube radius, R_(eff) is the lumen length of the capillary tubenot blocked by micro-bead flocculation, L is the capillary tube length,r the radius of the micro-beads, ϵ is the porosity of the micro-beadflocculation in the test micro-channel, and B is a constant that isintroduced to collected different terms under a single parameter foradding simplicity and clarity to the equation form. The model predictswith high accuracy (solid lines, FIG. 11) the curve shift with changesin capillary length over magnetic field region lengths. This modelprovided analytical guidance for designing the specifications of thedevice of the present invention.

(E) Microscopic Experimental Results

In order to observe the mechanism of magnetically-induced flocculation,microscopic studies were performed using an inverted microscope(Olympus, IX70, 20× magnification). This phenomenon was visualized usinga solution seeded with RBC-sized magnetic-responsive micro-beads, havinga diameter of 4 μm or 8 μm, in capillary tubes having a diameter of 50μm or 100 μm. FIG. 12 shows an example of flocculation 64 formed in acapillary channel 62 with glass walls 60, using a 2,000 micro-beads/μlanalog solution. Flocculation initially formed at the leading edge ofthe magnet 24 (corresponding to the greatest magnetic field gradient.)The size of the flocculation grew over the entire length of the magneticfield region. When the size of the flocculation covered the length ofthe magnetic field region, the flocculation behaved as a fluidized bed.Magnetic-responsive micro-beads downstream were released, while upstreamincoming magnetic-responsive micro-beads were added to the flocculation.

Example 2—pScreen™ Prototype Fabrication and Testing (A) Fabrication

Two sets of pScreen™ prototypes were fabricated: (1) a bench topprototype with multiple micro-channels for simultaneous testing ofvarious samples; and (2) a single-use, portable, lab-on-card device. ThepScreen™ bench-top prototype was described in the previous section. ThepScreen™ lab-on-card prototype was realized using standard microfluidicsfabrication techniques. In brief, the micro-channels were etched using alaser etcher system on a poly(ethylene terephthalate) glycol (Petg)substrate. The channels then were sealed using a matchingpoly(lactic-co-glycolic acid) (PLGA) top by thermal bonding using a hotpress. The magnetic field gradient was obtained by placing two smallmagnets in an N-S configuration underneath the test channel. ThepScreen™ lab-on-card prototype also was fabricated using injection castmolding in which the prototype was fabricated in poly(methylmethacrylate) (PMMA).

(B) Experimental Data for a Microfluidic Device for Detecting andQuantifying the Concentration of Magnetic Micro-Beads

Two pScreen™ prototypes were tested using a variety of fluids such as,without limitation, blood, blood-analogs, or PBS buffer solution withdifferent concentrations of surfactant. Tests were conducted usingmagnetic-responsive micro beads having a diameter of 4.1 μm or 8 μm.Sample concentrations between 100 micro-beads/μl and 200,000micro-beads/μl were used. The concentration of magnetic-responsivemicro-beads was determined by recording the level of the fluid on thecalibration and test graduated column scales. Each mark on the scalecorresponded to a given amount of fluid volume which flowed through themicro-channels. The relationship derived in Equation 8 was applied toconvert the recorded volumes in magnetic-responsive micro-beadconcentration. The analytic expression of the relationship between thevolumes Vo and Vm, specific for the tested prototypes, was derived bycomputing Equations (5) through (8), with

given in Equations (9) through (11). To be especially noted is the factthat all of the equations provided above do not include time as avariable. Hence, it is not necessary to monitor/read the device's resultat any specific time. The device reading at any time provides the sameread-out. FIG. 13 is a graph showing magnetic-responsive micro-beadconcentration measured using the pScreen™ device (y-axis) of the presentinvention versus magnetic-responsive micro-bead concentration measuredwith a standard hemocytometer (x-axis).

(C) Experimental Data for a pScreen™ Immunoassay

Several pScreen™ immunoassay prototypes were tested using buffersolutions containing various concentrations of mouse-IgG antibodyprepared by titration from a known concentration IgG standard. Theconcentration of mouse-IgG antibody ranged from 0.5 ng/ml to 100 ng/ml.Tests were conducted using magnetic-responsive micro-beads coated withanti-mouse IgG antibody and coating the surface of a reaction chamberwith anti-mouse IgG antibody. Sample volume ranged from 30 μl to 60 μl.The IgG antibody concentration was determined by recording the level ofthe fluid on the calibration and test graduated column scales. Each markon the scale corresponded to a given amount of fluid volume which flowedthrough the micro-channels. FIG. 14 is a graph showing the differencebetween the volume collected in the control column (Vo) and the volumecollected in the test column (Vm) (y-axis) versus the knownconcentration of IgG antibody (x-axis).

Examples 3 to 6 provide guidance regarding the detection andquantification of magnetic-responsive micro-beads and amount of analytein a liquid sample. Also provided are examples showing how to obtain aset of calibration curves for any embodiment by means of the claimedproxies herein, i.e., between the ratio {tilde under (O)}m/{tilde under(O)}o, the difference {tilde under (O)}o−{tilde under (O)}m, and theratio ({tilde under (O)}o−{tilde under (O)}m)^(p)/({tilde under(O)}m)^(q), and the number of magnetic-responsive micro-beads and theconcentration of analyte in a fluid. The examples provided show how toderive Equation 8 for any specific embodiment and provide workingexamples for a selected embodiment. Equation 8 is a direct result of theclaimed proxy between the ratio {tilde under (O)}m/{tilde under (O)}oand the number of magnetic-responsive micro-beads. Equation 8 convertsthe volume of liquid sample passing through the control micro-channel,Vo, and through the test micro-channel, Vm, into the magnetic-responsivemicro-bead concentration and the analyte concentration in a liquidsample.

Example 3—Proxy Relationships between Flow Rate and Micro-BeadConcentration

The present invention provides a method of detecting and quantifying theconcentration of magnetic-responsive micro-beads in a liquid sample, bycalculating the ratio {tilde under (O)}m/{tilde under (O)}o, thedifference {tilde under (O)}o−{tilde under (O)}m, and the ratio ({tildeunder (O)}o−{tilde under (O)}m)^(p)/({tilde under (O)}m)^(q), wherein pand q are derived through a calibration process, wherein the ratios{tilde under (O)}m/{tilde under (O)}o and ({tilde under (O)}o−{tildeunder (O)}m)^(p)/({tilde under (O)}m)^(q) are a proxy for the number ofmagnetic-responsive micro-beads in the liquid sample, and thenquantifying the concentration of magnetic-responsive micro-beads in theliquid sample.

Three methods are provided to show how to use these proxy relationshipsand how to derive in practice the number of magnetic-responsivemicro-beads in a liquid sample. The first method analytically derivesthese relationships from the physical dimensions of a specificembodiment. The second method provides a calibration process thatenables the calibration of any embodiment. The third method teaches howto use the volume of sample passing through the control micro-channel,Vo, and through the test micro-channel, Vm, to derive the concentrationof magnetic-responsive micro-beads.

Method 1. Relationship Between Normalized Flow Rate, {tilde under(O)}m/{tilde under (O)}o, and Number of Micro-Beads

The relationship between the normalized flow rate, {tilde under(O)}m/{tilde under (O)}o, and the number of micro-beads entering thecapillary tubes (i.e., micro-channels) is analytically derived for aspecific embodiment having the following dimensions and values:

-   -   Test micro-channel and calibration (control) micro-channel: each        with a circular cross-section with a radius of 50 μm; and a        length of 15 mm.    -   Micro-beads: radius of 2 μm    -   Porosity of the flocculation: 0.55    -   Pressure difference between the inlet and outlet of both        micro-channels: maintained constant at about 1 mmHg per cm.

The flow rate in the control micro-channel, {tilde under (O)}o, isdescribed by the Hagen-Poiseuille equation, and is equal to:

{tilde under (O)} _(o) =ΔP/R _(o),with: R _(o)=8μL/(πα⁴),   Equation(12)

where ΔP is the pressure drop across both micro-channels; Ro is themicro-channel's flow resistance; μ is the liquid sample viscosity; α isthe micro-channel's radius, and L is the micro-channel's length.

The flow rate in the test micro-channel, {tilde under (O)}m, is givenby:

{tilde under (O)} _(m) =ΔP/R _(m), with: R _(m)=8μ(L−λ)/(πα⁴)+μλ/(πα²k),   Equation (13)

where λ is longitudinal section of the length of the test micro-channeloccupied by the micro-bead flocculation; L−λ is the remaining length ofthe micro-channel flocculation-free; and k is the Darcy's constant. Thefirst term in Equation 13 is the Hagen-Poiseuille equation for theflocculation-free section of the test micro-channel, and the second termis Darcy's law for the section of the test micro-channel occupied by themicro-bead flocculation. Darcy's law is known by those skilled in theart to accurately describe flow of fluid through packed micro-beads andmicro-bead flocculation. In this example, the Reynolds' number is low(<2) (Reynolds' number depends on an embodiment's physical dimensionsand flow rates, and is a well know parameter that describes the physicalproperties of fluid in a flow), and thus the Darcy's constant takes theknown form:

k=r ²ϵ³/(15(1−ϵ)²),   Equation (14)

where ϵ is the porosity of the micro-bead flocculation in the testmicro-channel, i.e., the fraction of volume of the test micro-channelnot occupied by the micro-beads; and r is the micro-bead's radius.

Taking the ratio of Equation 13 over Equation 12, one gets:

${Q_{m}/Q_{o}} = {{R_{o}/R_{m}} = {\frac{8\mspace{14mu} \mu \; L}{\pi \; a^{4}}/\lbrack {{8\; {{\mu ( {L - \lambda} )}/( {\pi \; a^{4}} )}} + {\mu \; {\lambda/( {\pi \; a^{2}k} )}}} \rbrack}}$

By rearranging the terms in the above equation, one gets:

{tilde under (O)}_(m) /{tilde under (O)} _(o)1/[1+(λ/8L)·(α² /k−1)].  Equation (15)

It is noted that Equation 15 does not contain liquid sample viscosity.

The longitudinal section of the length of the test micro-channeloccupied by the micro-bead flocculation, λ, increases with the number ofmicro-beads, N, that have entered the test micro-channel and formed theflocculation under the effect of the magnetic field. These quantities, λand N, thus are related by the following identity:

λ=N/

, with:

≡¾α²(1−ϵ)/r ³   Equation (16)

where

a constant that is introduced to collect different terms under a singleparameter for adding simplicity and clarity to the equation.

By putting Equation 16 into Equation 15, one gets:

{tilde under (O)} _(m) /{tilde under (O)} _(o)=1/[1+(N /8L

)(α² /k−1)]  Equation (17)

or alternatively,

{tilde under (O)} _(m) /{tilde under (O)} _(o)=1/[1+

], where: α=(⅛L

)(α² /k−1)

Equation 17 provides the relationship that enables the measurement ofthe number of micro-beads by measuring the ratio {tilde under(O)}m/{tilde under (O)}o. Thus, in this example, Equation 17 becomes:

{tilde under (O)} _(m) /{tilde under (O)} _(o)=1/[1+2.63×10⁻⁶ ·N],  Equation (18)

In this example, Equation 18 is used as follows:

Given a fluid sample of unknown micro-bead concentration, for which theconcentration needs to be determined, the operator first runs the samplethrough the micro-channels as described above and measures periodicallythe flow-rates, {tilde under (O)}m and {tilde under (O)}o. Themeasurements can be done at any time while the liquid sample is flowinginto the micro-channels. Then, by putting the ratio, {tilde under(O)}m/{tilde under (O)}o, obtained from each measurement into Equation18, the operator gets the number of micro-beads that have entered thetest micro-channel, N, at the time the flow rate measurements have beentaken. To compute the micro-bead concentration, the operator divides thevalue of N by the total volume of fluid that passed through the testmicro-channel at the time flow rate measurements have been taken, whichalso is equal to the sum of the products of the test micro-channel flowrates by the length of the time interval between measurements. Theshorter the time interval between measurements, the more precise is theobtained value of the number of micro-beads and of the micro-beadconcentration. Table 1 provides the values obtained in this example:

TABLE 1 Time Qo Qm Vm Concentration (second) μl/sec μl/sec Qo/Qm N (μl)(micro-beads/μl) 54 0.01 0.0100 1.000 0.0E+00 8.5 0.0E+00 906 0.010.0094 0.940 4.0E+04 17.1 2.3E+03 1818 0.01 0.0093 0.934 4.4E+04 26.11.7E+03 2778 0.01 0.0089 0.888 7.9E+04 35.0 2.3E+03 3786 0.01 0.00850.845 1.1E+05 43.8 2.6E+03 4830 0.01 0.0082 0.816 1.4E+05 52.6 2.7E+035910 0.01 0.0079 0.789 1.7E+05 61.4 2.7E+03 7026 0.01 0.0076 0.7631.9E+05 70.3 2.8E+03 8190 0.01 0.0073 0.732 2.3E+05 79.1 2.9E+03 93900.01 0.0071 0.710 2.6E+05 87.8 2.9E+03 10614 0.01 0.0070 0.696 2.7E+0596.6 2.8E+03

Table 1 shows the time in seconds at which each measurement was taken;the experimental measurements of the flow rate, {tilde under (O)}o, inμl/second in the control micro-channel; and the flow rate, {tilde under(O)}m, in μl/second in the test micro-channel. (For the method to work,it is not important how the flow rates, {tilde under (O)}o and {tildeunder (O)}m, are measured. For example, the flow rates can be measuredby using commercially available flow meters placed in series with themicro-channels, or as described in detail above in Example 1. The ratio,{tilde under (O)}m/{tilde under (O)}o, then is computed and shown in thetable's fourth column, from which the number of micro-beads, N, isderived using Equation 18 and shown in the table's fifth column. Thevolume, Vm, is calculated by taking the sum of the products of the testmicro-channel flow rates by the length of the time interval betweenmeasurements, and is shown in the table's sixth column (alternatively,the volume, Vm, also can be measured as an independent variable, asdescribed above). Then, micro-bead concentration is derived by dividingthe N column by the Vm column, with the result shown in the table'sseventh column. It is noted that the concentration slightly variesbetween measurements, and thus to obtain a more precise value, theaverage of these measurements may be taken. In this example, themicro-bead concentration is equal to 2,570 micro-beads/μ1 of liquidsample.

Equation 18 provided herein was derived for this specific embodiment. Aperson skilled in the art, however, would understand that following thesame process, the relationship between the normalized flow rate and thenumber of micro-beads entering the test micro-channel can be derived forother embodiments by putting the appropriate value of the Darcy' sconstant in Equation 15. The Darcy's constant is well known by thoseskilled in the art to depend on the micro-channels' shape, geometry,physical dimensions, design, flow rates, as well as an embodiment'sReynolds number. By using the appropriate Darcy's constant, theinvention may be practiced for other embodiments having, for example,non-cylindrical micro-channels and/or non-spherical micro-beads, and/ornon-uniformly-sized micro-beads.

In addition, for other embodiments, in which micro-bead flocculationdoes not occupy the entire test micro-channel's cross-section, but onlyoccupies outer layers, the relationship between the normalized flowrate, {tilde under (O)}m/{tilde under (O)}o, and the number ofmicro-beads entering the capillary tubes is given in Equations 9 to 11.

Method 2. Calibration Process

Provided herein is the construction of a set of calibration curves whichprovide the relationship between the normalized flow rate, {tilde under(O)}m/{tilde under (O)}o, and the number of micro-beads entering thetest micro-channels. The method described herein allows one to avoid theneed to use complex analytical expressions, which for some embodimentsmay be too complex to be represented in a mathematical closed form suchas Equation 18.

For a given embodiment, the calibration process involves running aliquid solution containing a known concentration (the calibrationsolution) of magnetic-responsive micro-beads through the control andtest micro-channels, as shown in FIGS. 1 and 4. The flow rates, {tildeunder (O)}m and {tilde under (O)}o, then are measured at different timeintervals and recorded, and the total volume passing through the testmicro-channel also is recorded, as described above. The shorter the timeinterval between recordings, the more precise is the resultingcalibration curve. For each value of flow rate recorded, the number ofmicro-beads that have entered the test micro-channel and added to theflocculation is computed by taking the product between the volume, Vm,passing through the test micro-channel and the value of the knownconcentration of the micro-beads. Then, the ratio, {tilde under(O)}m/{tilde under (O)}o, for each measurement is plotted versus thenumber of micro-beads computed, as previously described. The process maybe repeated multiple times using calibration solutions with differentknown values of micro-bead concentration; the more times this process isrepeated, the more precise are the calibration curves. An example ofthese plots is shown in FIG. 11 for the embodiment described in theExample 1. Using this calibration curve, micro-bead number in any testedfluid of unknown concentration can be determined from the observedratio, {tilde under (O)}m/{tilde under (O)}o, by reading the horizontalvalue (the micro-bead number) on the calibration curve correspondingwith the vertical value (the observed ratio, {tilde under (O)}m/{tildeunder (O)}o) observed during the measurement. The concentration of thecalibration solutions are prepared to approximate the expected range ofconcentration of the liquid samples to be tested.

The same process can be repeated to derive calibration curves thatrelate the ratio, ({tilde under (O)}o−({tilde under (O)}m)^(p)/(({tildeunder (O)}m)^(q), with micro-bead concentration in a liquid sample. Inthis case, the parameters, p and q, are obtained as followed. A solutioncontaining a known concentration of micro-beads and of known volume ispassed through the micro-channels and the flow rates, {tilde under (O)}mand {tilde under (O)}o, are measured. Then, an identical solution ispassed through the micro-channels but with a higher flow rate, {tildeunder (O)}o. This process is repeated at least three times, with flowrates, {tilde under (O)}o, which cover the range of flow rates, {tildeunder (O)}o, for which the embodiment is expected to be used. Then,({tilde under (O)}o−{tilde under (O)}m)^(p)/({tilde under (O)}m)^(q),with p and q set equal to 1, are plotted for each solution versus thenumber of micro-beads entering the test micro-channel (determined aspreviously described) to obtain a set of calibration curves. Then, p andq are independently varied until each curve ({tilde under (O)}o−{tildeunder (O)}m)^(p)/({tilde under (O)}m)^(q) collapses on each other toform a single curve. A method to optimize p and q is to solve thefollowing set of equations.

[({tilde under (O)} _(o) −{tilde under (O)} _(m))^(p) /{tilde under (O)}_(m) ^(p)]_(n) ¹=[({tilde under (O)} _(o) −{tilde under (O)} _(m))^(p)/{tilde under (O)} _(m) ^(p)]_(n) ²

[({tilde under (O)} _(o) −{tilde under (O)} _(m))^(p) /{tilde under (O)}_(m) ^(p)]_(n) ¹=[({tilde under (O)} _(o) −{tilde under (O)} _(m))^(p)/{tilde under (O)} _(m) ^(p)]_(n) ³

where the superscript numbers 1, 2, and 3 refers to the calibrationcurve number, and the subscript N indicates that the values of {tildeunder (O)}o and {tilde under (O)}m in the square brackets are taken whenthe number of micro-beads entering the test micro-channel is equal to N.The value of N can be chosen arbitrarily, and the process may berepeated for different N values to obtain a more accurate value of p andq across the entire range of numbers of micro-beads.

Method 3. Conversion from Volume to Micro-Bead Concentration—Equation 8Working Example

Provided herein are two processes which may be used for any embodimentin order to convert the volume of sample passing through the controlmicro-channel, Vo, and the test micro-channel, Vm, intomagnetic-responsive micro-bead concentration and concentration ofanalyte in a liquid sample. The first process analytically derives theserelationships from the physical dimensions of the embodiment. The secondprocess provides a calibration method that enables calibration of anyembodiment and shows the use of the embodiment in practice.

Process A

Provided is a working example showing how Equations 5 through 8 enableone to derive the concentration of micro-beads in a liquid from theratio {tilde under (O)}o/{tilde under (O)}m, using the same dimensionsprovided above in Method 1 (spherical micro-beads with a radius of 2 μm,micro-channels having a length of 15 mm and a radius of 50 μm, and{tilde under (O)}o equal to 0.0 μl/second), and for which the ratio,{tilde under (O)}m/{tilde under (O)}o, versus the number of micro-beadsis provided by Equation 18. By putting Equation 18 into Equation 5, onegets:

$\begin{matrix}{N_{0} = {\int\limits_{0}^{Nm}\frac{{dN}^{\prime}}{( {1/( {1 + {( {2.63 \times 10^{- 6}} ) \cdot N^{\prime}}} )} )}}} & {{Equation}\mspace{14mu} (19)}\end{matrix}$

where No is the number of magnetic responsive micro-beads that arepassing through the control micro-channel, Nm is the number of magneticresponsive micro-beads that are passing through the test micro-channel,and the prime symbol inside the integral, N′, indicates, according tostandard convention, that the integral operation is computed on an Nvariable. Equation 19 is the specific example, for a specificembodiment, of generic Equation 6. By solving the integral, one gets:

$\begin{matrix}{N_{o} = {N_{m} + {\frac{2.63 \times 10^{- 6}}{2}( N_{m} )^{2}}}} & {{Equation}\mspace{14mu} (20)}\end{matrix}$

Since No is equal to the product of the magnetic responsive micro-beadconcentration, p, times the volume of sample passing through the controlmicro-channel, Vo; and since Nm is equal to the product of the magneticresponsive micro-bead concentration, ρ, times the volume of samplepassing through the test micro-channel, Vm, Equation 20 can be rewrittenas:

$\begin{matrix}{{{{V_{o}\rho} = {{V_{m}\rho} + {\frac{2.63 \times 10^{- 6}}{2}( {\rho \; V_{m}} )^{2}}}};}{{{{or}\text{:}\mspace{14mu} V_{o}} - V_{m}} = {\frac{2.63 \times 10^{- 6}}{2}( V_{m} )^{2}\rho}}} & {{Equation}\mspace{14mu} (21)}\end{matrix}$

Equation 21 can easily be rearranged as follows:

$\begin{matrix}{\rho = {( {V_{0} - V_{m}} )/( {\frac{2.63 \times 10^{- 6}}{2}V_{m}^{2}} )}} & {{Equation}\mspace{14mu} (22)}\end{matrix}$

Equation 22 provides a simple expression that can be used to derive theconcentration of micro-beads for any liquid sample with unknownconcentration of micro-beads from the measurements of the volumes, Voand Vm, in this working example. In other words, Equation 22 is aspecific example of generic Equation 8. For example, a liquid samplecontaining an unknown concentration of magnetic-responsive micro-beadsis run through the control and test micro-channels, as described above.To obtain the value of this unknown concentration, the volume of thefluid passing through the control micro-channel, Vo, and the volume offluid passing through the test micro-channel, Vm, are recorded, asdescribed above. Then these values, Vo and Vm, are put into Equation 22to derive the concentration of magnetic-responsive micro-beads.

Process B

Provided herein is an example showing how to construct calibrationcurves for the relationship between volumes, Vo and Vm, andconcentration of micro-beads. The process allows one to avoid the needto use complex analytical expressions, which for some embodiments may betoo complex to be represented in a simple closed form such as Equation22. For a given embodiment, a series of solutions of known volume andcontaining known concentrations of magnetic-responsive micro-beads arepassed through the control and test micro-channels, as described above.The concentrations of these test (calibration) solutions are chosen tospan the entire range of micro-bead concentrations in the samples to betested. The volumes, Vo and Vm, then are measured. For the lab-on-cardmicrofluidic pScreen™ magnetic-responsive micro-bead concentrationcounter embodiment, the volumes, Vm and Vo, can be measured by readingthe scale on the control column (Co) and the test column (Cm). Once Vmand Vo have been recorded for all solutions, the ratios, (Vo−Vm)/(Vm)²,are plotted as a function of micro-bead concentration. This plotprovides a calibration curve that enables the invention to be reduced topractice. Using this calibration curve, the micro-bead concentration ofan unknown fluid can be determined from the observed difference, ratio(Vo−Vm)/(Vm)², by reading the horizontal value (micro-beadconcentration) on the calibration curve corresponding with the y value(Vo−Vm)/(Vm)² observed during the test, as shown in FIG. 15. Forexample, as shown in FIG. 15, if the measurement of Vm and Vo for asolution with an unknown micro-bead concentration gives a ratio of(Vo−Vm)/(Vm)² equal to about 0.02, then the micro-bead concentrationwould be about 7500 micro-beads/μl. The calibration curve shown in FIG.15 was obtained using the embodiment shown in FIG. 2, in which themicro-channels' radii is 46 μm, the length of the micro-channels is 15mm, and the radius of the micro-beads is 2.35 μm.

Example 4—Proxy Relationships Between Flow Rate and Concentration ofAnalyte in a Liquid Sample

The present invention provides a method of detecting and quantifying theconcentration of magnetic-responsive micro-beads in a liquid sample, bycalculating the ratio {tilde under (O)}m/{tilde under (O)}o, thedifference {tilde under (O)}o−{tilde under (O)}m, and the ratio ({tildeunder (O)}o−{tilde under (O)}m)^(p)/({tilde under (O)}m)^(q), whereinthe ratios {tilde under (O)}m/{tilde under (O)}o and ({tilde under(O)}o−{tilde under (O)}m)^(p)/({tilde under (O)}m)^(q) are a proxy forthe number of magnetic-responsive micro-beads in the liquid sample,which is a proxy for the concentration of analyte in the liquid sample,and then quantifying the concentration of analyte in the liquid sample,as described below.

As described above, the concentration of magnetic-responsive micro-beadsexiting the reaction chamber and entering the two micro-channels isproportional to the concentration of the analyte in the liquid sample.As described herein, the concentrations of these micro-beads aredetectable and quantifiable by either using the ratio {tilde under(O)}m/{tilde under (O)}o or the values Vo and Vm. Thus, in order toderive the concentration of analyte in a liquid sample, the relationshipbetween the concentration of analyte and the concentration ofmicro-beads exiting the reaction chamber needs to be calibrated. Twomethods are provided, as follows:

Method 1

The first method consists of running a sample with a known concentrationof analyte through the reaction chamber and then counting the number ofmicro-beads binding to the reaction chamber's surface, as shown in FIGS.8 and 9. The number of micro-beads exiting the reaction chamber is equalto the total number of micro-beads deposited in the reaction chamber (asdescribed above) minus the number of micro-beads binding to the reactionchamber's surface. The process is repeated multiple times using a seriesof calibration samples obtained from the initial sample by titration.Then the concentration of micro-beads exiting the reaction chamberobtained for each calibration test is plotted as a function of theanalyte concentration. In practice, the sample with unknownconcentration of analyte is processed as described above andschematically shown in FIGS. 1 and 7. The concentration of micro-beadsentering the micro-channels, i.e., the concentration of micro-beadsexiting the reaction chamber, is derived as described above. Then, usingthe calibration curve derived and described above, the concentration ofmicro-beads is converted into the concentration of analyte.

Method 2

The second method enables the user to derive the calibration curve forany embodiment without counting the number of micro-beads binding to thereaction chamber's surface. To calibrate for any embodiment, a series ofsolutions of known volume, containing known concentrations of targetanalyte, are run through the device as described above. Theconcentrations of these test (calibration) solutions are chosen to spanthe entire range of expected analyte concentrations of the sample to betested. The difference, Vo−Vm, between the control micro-channel volume,Vo, and the test micro-channel volume, Vm, is measured for each testedcalibration sample and plotted as a function of the calibration sample'sanalyte concentration. Using this plot (calibration curve), the targetanalyte in any tested fluid can be determined from the observeddifference in volume, Vo−Vm, by reading the horizontal value (analyteconcentration) on the calibration curve corresponding with the verticalvalue (the Vo−Vm difference) observed in the test.

Example 5—Working Example—Quantification of Concentration of Micro-Beadsin a Liquid Sample

Quantification of the concentration of micro-beads in a liquid sample isperformed by using the pScreen™ immunoassay shown in FIG. 2, and havingthe dimensions as described above in Method 3, Process A. To measure theconcentration of micro-beads, the user performs the following steps:

Step 1: Take 50 μl of sample solution and place it on the device inlet.

Step 2: Wait for the sample to flow into the device columns, typicallyless than 20 minutes.

Step 3: Read the values of the fluid volumes, Vo and Vm, filling up eachcolumn using the scale on the side of the column.

Step 4: Compute the ratio (Vo−Vm)/(Vm)².

Step 5: The concentration of micro-beads is equal to ratio (Vo−Vm)/(Vm)²divided by 1.3×10⁻⁶.

Example 6—Working Example—Quantification of Analyte Concentration in aLiquid Sample

Quantification of the concentration of an analyte in a liquid sample isperformed by using the pScreen™ immunoassay shown in FIG. 3, and thecalibration curve shown in FIG. 15. To measure the concentration of theanalyte, the user performs the following steps:

Step 1: Take 50 μl of sample solution and place it on the device inlet.

Step 2: Wait for the sample to flow into the device columns, typicallyless than 20 minutes.

Step 3: Read the values of the fluid volumes, Vo and Vm, filling up eachcolumn using the scale on the side of the column.

Step 4: Compute the ratio (Vo−Vm)/(Vm)².

Step 5: Read, on the calibration plot, the analyte concentration on thehorizontal axis of the calibration curve corresponding to the value ofthe (Vo−Vm)/(Vm)² ratio on the vertical axis.

Example 7—System of Valves

By introducing a system of valves in the present invention, theincubation time can be adjusted by varying the length of the delaymicro-channel. For any practical purpose, the incubation time is equalto the delay time, which is the time it takes fluid to travel the entirelength of the delay micro-channel from the primary flow splitter 74 tothe connector 67 (shown in FIG. 21). This time depends on severalfactors, including fluid type, delay micro-channel contact angle, delaymicro-channel width, depth, and length, and reaction chamber and inletmanifold's fluid dynamic resistance. The delay time is the time t thatsatisfies the following equation:

$\begin{matrix}{\mspace{79mu} {{l = {\text{?}( {\sqrt{R_{0}^{2} + {2D\text{?}}} - \sqrt{R_{0}^{2}}} )}},{\text{?}\text{indicates text missing or illegible when filed}}}} & {{Eq}.\mspace{11mu} 23}\end{matrix}$

where the following definitions apply:

$\mspace{20mu} {\text{?} = \frac{12\mspace{14mu} \mu}{{wh}^{3}( {1 - {0.63{h/w}}} )}}$$\mspace{20mu} {D = \frac{\gamma ( {{\cos ( \theta_{c} )}/r_{c}} )}{wh}}$?indicates text missing or illegible when filed

and, where: t,l(t),w,h,μ,γ,θ_(c)r_(c), R₀ are, respectively, the time ittakes for the fluid to travel the entire length of the delaymicro-channel; the length, width and depth of the delay micro-channel,the viscosity of the liquid sample, the surface energy of the liquidsample, the mean of the contact angle of the liquid sample with thewalls of the delay micro-channel, the hydraulic radius of the delaymicro-channel, and the sum of the flow resistances of the inletmanifold, reaction chamber, and outlet manifold micro-channels. Becauseof the complex physical behavior at the fluid meniscus, especially witha non-Newtonian liquid sample such as whole blood, the above equationprovides only an approximate solution, and the desired value needs to bedetermined experimentally in each case.

As a specific example, a series of experimental tests were performed andresults were obtained using whole blood, a device made of PMMA, sealedwith a hydrophilic tape, having a contact angle between the fluidmeniscus and the PMMA surface of between about 70 degrees and about 80degrees, and between the fluid meniscus and the hydrophilic tape ofbetween about 10 degrees and about 34 degrees.

In the first test, the delay time was determined experimentally to be 45sec and 64 sec for a delay micro-channel having a length of 40mm and60mm, respectively, and a width and depth of 0.25 mm and 0.25 mm,respectively.

In a second test, the delay time was 152 sec, for a delay micro-channellength, width and depth of 57 mm, 0.25 mm, and 0.1 mm, respectively.

In a third test, the delay time was 180 sec, for a delay micro-channelhaving a first section of length, width and depth of 30 mm, 0.2 mm, and0.2 mm, respectively, a second section of length, width and depth of 30mm, 0.1 mm, and 0.1 mm, respectively, and a third section of length,width and depth of 2 mm, 0.35 mm, and 0.25 mm, respectively.

Also provided is a method to determine the size of the appendixmicro-channels. As shown in FIG. 22, and as previously discussed, thepassive valve 66 substantially stops the fluid flow rate to a negligiblerate when the Pi>Pv, where Pi is the atmospheric pressure and Pv is thepressure inside the fluid created by the curvature of the meniscus. Ifthe sealing layer is hydrophobic, Pv is always larger than Pi, but ifthe sealing layer is hydrophilic, then Pv may be smaller than Pi causingthe valve to fail (FIG. 22 shows the case of a hydrophilic surface whichcreates a concave meniscus). By applying well-known hydrodynamic laws offlow in capillary circuits, it is shown that:

$\begin{matrix}{{{{Pi} - {Pv}} = {( {{Pi} - {Pl}} )\frac{R_{0}}{R_{o} + R_{l}}}},} & {{Eq}.\mspace{11mu} 24}\end{matrix}$

where R0 and R1 are, respectively, the flow resistance of the reactionchamber and of the delay micro-channel (which can be calculated by usingwell-known formulae), and

Pl=γcos(θ_(c))/r _(c),

Pν=2γcos(θ_(c))/w_(v)+γcod(θ_(c))/h _(v)−γcod(φ_(c))/w_(v)

where φ_(c), w_(v), h_(v) is the contact angle of the fluid with thesealing layer, and the width and depth of the valve. The size (width,depth, and length) of the appendix micro-channels are chosen such that(Pi−Po)>(Pi−Pv), where (Pi−Pv) is given by Eq. 24, and Po is thepressure at the secondary and primary flow split, and thus provides apressure gradient at the valve that prevents flow through the valve. Thevalue of Po is approximately given by the ratio of the flow rate of thefluid in the appendix micro-channel over the R1. The same applies forthe computation of the delay micro-channel's size.

Example 8—Conical Sample Inlet

A conical shape inlet with a top opening of 10 mm, a bottom inlet of 1mm, and height of 5 mm, was fabricated. The inner surface of the inletwas coated with a super-hydrophobic layer with a contact angle of 150degrees. The inlet then is attached to an embodiment of the microfluidicdevice described above in paragraph [0049], human plasma was placedinside the inlet, and then the flow rate of the liquid sample, when thefluid reached the graduated columns, was measured and compared to themicrofluidic device without a conical inlet. The flow rate without aconical inlet was 0.016 μl/sec, whereas with the conical inlet the flowrate increased to 0.022 μl/sec. Thus, the microfluidic device having aconical inlet in accordance with an embodiment of the present inventionresulted in an increased flow rate of about 37%.

Example 9—Micro-Bead Deposition using a Nano-Dispenser to DistributeMicro-Beads, with Addition of Sucrose and Micro-Bead Layering to PromoteRelease

Depositing micro-beads on the surface of a reaction chamber allows forprecise control of micro-bead placement and distribution. As shown inFIG. 23, micro-bead drops 83 are deposited in arrays 92 on the surfaceof a reaction chamber 34. The arrays 92 can be positioned based onvarious factors including, but not limited to, different reactionchamber sizes and geometries, micro-bead sizes and concentrations, andantibody affinities. FIG. 23 shows a particular array pattern. Each spotin the array is comprised of several layers of nano-drops. When a singlelayer of drops containing micro-beads suspended in a 10% sucrosesolution is deposited in 5-10 nl drops and dried in a dessicator, lessthan 50% of the micro-beads are released, i.e., rehydrated, when aliquid sample is added the reaction chamber. When two layers of thesedrops are deposited on top of each other, greater than 60% of themicro-beads are released upon addition of a liquid sample. When four ormore layers of drops are deposited on top of one another, greater than90% of micro-beads are released when the sample is added.

Example 10—Alteration of Micro-Bead Density to Achieve Neutral Buoyancyand to Prevent Micro-Bead Settling during Micro-Bead Deposition

Experiments were conducted to assess the settling of micro-beads insolutions with densities adjusted to achieve neutral buoyancy, and thusprevent micro-beads from settling and the concentration of themicro-bead solution from changing during dispensing. Keeping themicro-beads in a well-mixed suspension is essential to achieveuniformity in micro-bead distribution within and between reactionchambers.

In an exemplary experiment, suspending micro-beads in a 30% v/v glycerolsolution (35% w/v) prevented the micro-beads from settling. Three dropswere dispensed every 10 minutes, and micro-beads in each drop werecounted. As shown in FIG. 24, micro-bead count in each drop remainedconstant over a two-hour period. No significant change in micro-beaduniformity was observed in a test tube or in a nano-dispenser reservoirand tubing over the two hour test period. In contrast, in the controlwhere the micro-beads were suspended in 10% sucrose dissolved in water,which has a density lower than the micro-beads, the micro-bead count ineach drop decreased by greater than 95% in two hours.

While the addition of glycerol is effective in preventing settling andpromoting uniformity in micro-bead deposition, the increased viscositydue to the glycerol may adversely affect the immunoassay. Thus, heavywater (deuterium oxide) was tested for neutral buoyancy suspension ofmicro-beads for deposition of nano-drops. Micro-beads were suspended ina solution of heavy water (deuterium oxide) mixed with water to achievea density that matched that of the micro-beads. Sucrose (10%) was addedto promote micro-bead release from the reaction chamber surface duringthe immunoassay. A solution containing between 80-90% heavy waterallowed the micro-beads to remain in solution without settling. Nosignificant change in micro-bead uniformity was observed in a test tubeor in a nano-dispenser reservoir and tubing over a two hour test period.Micro-bead count in each drop remained relatively constant over the hourperiod (shown in FIG. 24) with the count at two hours being less than 5%lower than the initial count (compared to a greater than 95% reductionin count in the control). Suspension of micro-beads in a solutioncontaining heavy water (deuterium oxide) allows for uniformity indeposition of micro-beads without affecting the assay performance, sinceheavy water has the same viscosity and chemical properties as water.

Example 11—Hydrophilic Coating of Micro-Channels to increase Flow Rateand Decrease Assay Time

All micro-channels were coated using a protein-free blocking solution.This coating creates a hydrophilic film on the surface that decreasesthe fluid contact angle, thus increasing flow rate and decreasing assaytimes. This coating decreased the contact angle of the PMMA from about74° to less than about 10°. A liquid sample was placed in themicro-channels through the liquid sample inlet and flow rate wasrecorded. This coating caused an average increase in flow rate of plasmaby 70% (e.g., from 0.016 μl/sec to 0.028 μl/s) and in whole blood by afactor of two (e.g., from 0.005 to 0.01 μl/s) compared to flow of thesame sample in uncoated channels of the same dimensions. This increasein flow rate allows the assay time to be reduced by as much as 50%.

Example 12—Flow Rate Variation of Human Whole Blood through the pScreen™Immunoassay Device

Flow rate of human whole blood, anticoagulated with EDTA, was measuredin a pScreen™ immunoassay device (shown in FIG. 18).

The device had the following dimensions:

Inlet manifold micro-channel: width, depth and length between 100 μm-300μm, 100 μm-300 μm, and 5 mm-15 mm, respectively.

Reaction chambers: width, depth, and length between 120 μm-200 μm, 0.5mm-3 mm, and 20-40 mm, respectively.

Outlet manifold micro-channels: width, depth and length between 50μm-300 μm, 50 μm-300 μm, and 5 mm-20 mm, respectively.

Delay micro-channel and appendix micro-channels: width, depth, andlength between 50 μm-200 μm, 50 μm-200 μm, and 30 mm-150 mm,respectively.

Test and calibration micro-channels: width, depth and length between 50μm-200 μm, 50 μm-200 μm, and 5 mm-20 mm, respectively.

The mean flow rate when the fluid meniscus is in the reaction chamber,delay micro-channel or appendix micro-channels, connector micro-channel,and test and calibration micro-channels is between 0.15 μl/sec to 0.25μl/sec; 0.035 μl/sec to 0.005 μl/sec; 0.12 μl/sec to 0.08 μl/sec, and0.06 μl/sec to 0.3 μl/sec, respectively.

Example 13—Assay Calibration

Experiments were conducted to assess the sensitivity and dynamic rangeof the pScreen microfluidic device in quantifying the concentration ofDNA fragments in solution.

Buffer 1 was prepared at a final concentration of 1.2 Mole/Liter AceticAcid (CH₃COOH), 1 Mole/Liter sodium chloride (NaCl), and 20% v/v PEG400, and containing nonmagnetic-responsive silica micro-beads, with sizeranging from 1 μm to 9 μm, in a concentration of 30 μg/μL, andmagnetic-responsive silica beads in concentration of 1.3 μg/μl. In placeof Buffer 2, Monarch® DNA Cleanup Binding Buffer was used, which is achaotropic salt-based binding buffer containing isopropyl alcohol,guanidine hydrochloride, and sodium Perchlorate.

Samples were prepared my mixing de-ionized water spiked with a knownconcentration of DNA fragments. Different lengths of DNA fragments wereselected ranging from 150 base pairs (BPs) to over 10,000 BPs. First, 5μL of DNA-containing sample was pipetted into a clean 1.5 mLmicrocentrifuge tube. Subsequently, 25 μl of Buffer 1 was added to thesample in the 1.5 mL microcentrifuge and mixed by pipetting up and downbetween 1 and 5 times. The solution was then incubated for 1 minute.Then, 100 μL of Solution 2 was added, and while holding the tube usingthe thumb and forefinger of one hand, the 1.5 mL microcentrifuge tubewas flicked vigorously for 30 seconds using the forefinger of the otherhand. Finally, the solution contained in the 1.5 mL microcentrifuge tube(i.e., the sample mixed with Buffer 1 and 2, as above described) waspipetted up and down between 1 and 10 times and transferred into theinlet of the microfluidic device. The test is completed when the liquidin the control column reaches the passive valve, and the flow in themicrofluidic card stops. The location where the liquid in the testcolumn stops was recorded, and its distance to the passive valve,Delta-L, plotted against the Scaled Concentration (ng/μl) of the DNAfragments (see FIG. 27). The Scale Concentration is defined as theactual DNA fragment concentration divided by a calibration BP-dependentconstant, Cal. The calibration constant, Cal, is experimentallydetermined so that each Delta-L vs Scale Concentration, for all BPs,collapses as shown in FIG. 27 on the same master curve.

Sample volume used for this set of experiments varied between 1 μl and 5ul. The dynamic range of the assay with 5 ul input is approx. 5 ng/uL to30 ng/uL. This can be extended further up to 150 ng/ul by loweringsample volume to 1 ul. Accordingly, as shown in FIG. 27, the overalldynamic range is approximately 5 ng/ul to 150 ng/ul, thus overlappingwith most molecular biology methods. Samples with DNA concentrationsexceeding 150 ng/ul can be diluted prior to measurement to avoidsaturation, using water or other buffer DNA compatible buffers (e.g. TEbuffer).

Example 14—Assay Validation

The assay was validated using multiple DNA sample types, includinghighly purified NoLimits DNA Fragments (ThermoFisher), fragmentedgenomic DNA (ThermoFisher), and products generated from standardamplification methods (PCR). FIG. 28 summarizes the results fromexperiments designed to evaluate the correlation between DNAconcentrations of multiple samples measured using the microfluidicdevice and the same samples analyzed with UV spectrometry (purified DNA)or gel electrophoresis in conjunction with densitometry (unpurified

DNA). The validation study was performed using the same Buffer 1 andBuffer 2 as described in Example 13.

Each sample was prepared using Buffer 1 and Buffer 2, and the sample runon the pScreen assay as described above. The resulting Delta-Ls wererecorded for each sample, and the scaled concentration of the DNAfragments subsequently was calculated from FIG. 27, by reading the valueof the Delta-L on the y-axis of FIG. 27, and recording the correspondentvalue on the x-axis, i.e., the scaled concentration. Then, using the Calconstant as derived in Example 13, the concentration of DNA fragmentswas computed by multiplying the scaled concentration by the Calconstant. This process can also be accomplished using an algorithmcalculator or, alternatively, using a manual calibration table.

FIG. 28 shows the value of Delta-L as a function of the actual DNAfragment concentration for various concentrations of DNA fragments andfragment size from 150 BPs to over 10,000 BPs. FIG. 28 shows that a highdegree of correlation was observed between the concentration of DNAfragments measured by the pScreen assay as described above and theactual (true) DNA concentration. The graph shows a correlation of 1, anda R² =0.97. The standard deviation is about 10%. No significant bias inthe results obtained with the pScreen assay was noted with regard toeither DNA fragment length or concentration. As it is apparent byexamining FIG. 28, one key aspect that makes the microfluidic device anattractive option for DNA quantification is its insensitivity to theinterference stemming from the presence of typical PCR reactioncomponents, thereby eliminating the need for separate purificationsteps. Taken together, the collected data support the notion that themethodology described above combines specificity, sensitivity, speed andcost-effectiveness in a simple and user-friendly procedure which shouldfind multiple applications in the field of molecular biology.

It will be appreciated by those skilled in the art that changes could bemade to the embodiments described above without departing from the broadinventive concept thereof. It is understood, therefore, that thisinvention is not limited to the particular embodiments disclosed, but itis intended to cover modifications that are within the spirit and scopeof the invention, as defined by the appended claims.

What is claimed is:
 1. A microfluid device for quantifying theconcentration of DNA fragments, comprising: a sample inlet defined by anopening for accepting a liquid sample; a sealing layer; a reactionchamber; a micro-channel splitter; a calibration micro-channel; a testmicro-channel a control graduated column; a test graduated column; aventing hole located at the end of each of the graduated columns that isplaced through the sealing layer; and a magnetic field gradient adjacentto the test micro-channel.
 2. The microfluid device of claim 1, whereinthe liquid sample containing a mixture of magnetic-responsive silicamicro-beads, nonmagnetic-responsive silica micro-beads, a chaotropicbuffer, and an aliquot of a DNA fragment sample is placed on the inletof the microfluidic device.
 3. The microfluid device of claim 1, whereinthe presence of DNA fragments triggers the formation of largemagnetic-responsive clusters, said magnetic-responsive clustersaggregating in proximity of the magnetic field gradient to produce alocalized restriction, said restriction retarding flow of the liquidsample in the test column.
 4. The microfluid device of claim 2, whereinthe concentration of nonmagnetic-responsive silica micro-beads rangesfrom between 10 μg/μl and 60 μg/μL in the liquid sample.
 5. Themicrofluid device of claim 2, wherein the concentration ofmagnetic-responsive silica micro-beads ranges from between 1.0 μg/μl and3.0 μg/μL in the liquid sample.
 6. The microfluid device of claim 1,wherein the control graduated column and the test graduated column areconnected to one another via a joining channel, and a venting holelocated at the end of the control column is placed through the sealinglayer to stop the flow in the test channel when the liquid flowing inthe control column fills to a desired volume.
 7. A microfluid passivevalve layout to stop the flow of liquid in a primary micro-channel whena control micro-channel fills to a desired volume, comprising: a primarymicro-channel; a control micro-channel; a joining channel connecting thecontrol micro-channel and the primary micro-channel; a venting hole inthe joining channel, said venting hole placed through the sealing layer,said venting hole placed such that the volume of liquid filling thecontrol micro-channel is a desired amount.